Light-adjustable hydrogel and bioanalogic intraocular lens

ABSTRACT

A bioanalogic implantable ophthalmic lens (“BIOL”) capable of replacing the natural crystalline lens (NCL) in its various essential functions after the NCL having been removed and BIOL implanted into the posterior eye chamber and placed into the capsular bag vacated from the NCL. At least the posterior surface of the lens has a convex shape and is made from a transparent flexible hydrogel material. At least the anterior and posterior optical surfaces are defined by rotation of one or more conic sections along the main optical axis and the surfaces defined by the rotation will include a plane perpendicular to the axis and conical surface symmetrical by the axis. A hydrogel implantable ophthalmic lens whose optical parameters can be optimized and/or customized by a controlled absorption of electromagnetic radiation resulting in a change of the refractive index of the irradiated hydrogel.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application is a Continuation of U.S. application Ser. No.15/190,715, filed Jun. 23, 2016, which is a Continuation-in-Part of U.S.application Ser. No. 14/760,868, filed on Jul. 14, 2015, which is theU.S. National Phase of International Application No. PCT/IB13/60869,filed on Dec. 12, 2013, which claims the benefit of priority of U.S.Provisional Application No. 61/752,685, filed on Jan. 15, 2013. Theentire contents of all of the above applications are incorporated hereinby reference.

FIELD OF THE INVENTION

This invention relates to a hydrogel implantable ophthalmic lens whoseoptical parameters can be optimized and/or customized by a controlledabsorption of electromagnetic radiation, such as laser radiation in thevisible and/or near-IR region, and more particularly radiation emittedin pulses shorter than one nanosecond (so called Femtosecond Laser, FSL)resulting in a change of the refractive index of the irradiatedhydrogel.

BACKGROUND OF THE INVENTION

Intra ocular lenses (IOLs) are surgically implantable lenses whichreplace or supplement optical function of the NCL. So called “posteriorchamber intraocular lens”, or PC IOLs, replace the NCL in the case ofcataract or, more recently, in the case of presbyopia by so called“clear lens exchange”, or CLE. Other implantable lenses are placed intothe anterior chamber of the eye (AC IOLs), into the cornea (corneal orintrastromal implants) or between the NCL and iris (so called“implantable contact lens” or ICL). So far, most of these IOLs weredesigned to replace or to supplement the basic optical function of theNCL only. It should be appreciated that an NCL in a human eye, depictedin the FIG. 1, is a complicated structure with several functions. Themain eye parts include the cornea 101; the iris 102; the NCL 103; theposterior capsule 104; the cilliary muscle 105; the zonules 106; thevitreous body 107; and the retina 108.

The basic optical function of the NCL 103 consists in helping the cornea101 to focus the incoming light so that a distant object can beprojected on the retina 108. The other important optical function isaccommodation—adjustment of optical power of the lens in such a way thatobjects at various distances can be projected onto the retina 108. Thereare several theories explaining the accommodation mechanism. See forexample L. Werner et al, Physiology of Accommodation and Presbyopia,ARQ. BRAS. OFTALMOL. 63(6), DEZEMBRO/2000-503.

The most firmly established theory is von Helmholtz theory explainingthat, referring to the FIG. 1, relaxed cilliary muscle 105 causestension in the zonules 106 that pull the lens 103 periphery outward tokeep the NCL 103 in its deformed (flattened) shape that provides a lowerrefractive power suitable for distant vision. Focusing on a near objectis caused by tension in the cilliary muscle 105 that relaxes the zonules106 and allows the NCL 103 to obtain its “natural” configuration with asmaller diameter, larger central thickness and smaller radii ofcurvature on both anterior and posterior surfaces. This increases theNCL's refractive power and allows for projection of the image of nearobjects on the retina 108.

Most of the common intraocular lenses have spherical surfaces that canbe manufactured rather readily. It has been assumed for some time thatthe NCL 103 is essentially spherical. However, a spherical lens is notexactly monofocal, instead it demonstrates so called “sphericalaberration” wherein rays incoming through the center are bent into afocal point that is slightly further from the lens than rays incomingthrough the lens periphery. Therefore, a spherical lens is somewhat morerefractive in its periphery than in its center. This change iscontinuous: such a lens does not have a single focal point, but manyfocal points in a short interval of distances (focal range) between thelongest and shortest focal distance. In other words, a spherical lens isnegatively polyfocal (its focal distance decreases from the center tothe periphery). Lenses with elliptical rather than spherical surfaces(such as surfaces created by solidification of a static liquid meniscus)have even more distinct spherical aberration and are, therefore, evenmore negatively polyfocal than spherical lens.

Some artificial intraocular lenses include hyperbolic surfaces alongsidewith other surfaces of second order, such as spheric or even ellipticsurfaces that have negative polyfocality and very opposite opticaleffect. More importantly, the prior art generally combines second order(or conic section) surfaces with meniscoid surfaces that are poorlydefined and merely approximate second order surfaces with positivespherical aberration (although never surfaces with hyperbolicaberration).

For example, Wichterle in U.S. Pat. No. 4,971,732 claims the meniscoidsurfaces to approximate a flat ellipsoid while Stoy in U.S. Pat. No.5,674,283 considers meniscoid surfaces an approximation of a sphericalsurface, both having negative polyfocality. A combination of surfaceswith positive and negative polyfocality diminishes or negates advantagesof the former.

Furthermore, Wichterle '732 describes a manufacturing method of theintraocular lens where a monomer solidifies in an open mold, one(posterior) side of the lens having the shape of the mold cavity whilethe anterior side has a shape of a solidified liquid meniscus(presumably approximating a flat ellipsoid shape with negativepolyfocality, being somewhere between purely spherical and purelyellipsoid surface). The mold cavity has the shape of a second ordersurface that may include a hyperbolic surface. One can note that each ofthe optical surfaces is created differently—one by solidification of apolymer precursor against a solid surface while the other bysolidification on the liquid-gas interface. It is known to those skilledin the art that the surface quality of the two optical surfaces formedunder such different circumstances may differ profoundly in both opticaland biological respects.

Wichterle in U.S. Pat. No. 4,846,832 describes another manufacturingmethod of the intraocular lens where the posterior side of the lens hasthe shape of the solidified liquid meniscus (presumably approximating aflat ellipsoid shape with negative polyfocality) while the anterior sideis formed as an imprint of the solid mold shaped as a second ordersurface that may implicitly include also a hyperbolic surface. Again, wecan note that each of the optical surfaces is created differently—one bysolidification of a polymer precursor against a solid surface while theother by solidification on the liquid-gas interface.

Stoy '283 discloses modifying the method described by Wichterle '732using a two part mold, one part being similar to the Wichterle's moldwhile the other being used to form a modified meniscoid of a smallerdiameter on the anterior lens surface. The meniscoid optical surface isof the same character as the meniscoid resulting from Wichterle '732,albeit of a smaller diameter and, therefore, probably closer to aspherical surface than an ellipsoid surface. In any case, such a surfacehas negative polyfocality. The posterior side is formed as an imprint ofthe solid mold shaped as a second order surface that may include ahyperbolic surface while the other optical surface is formed bysolidification of the liquid polymer precursor on the liquid-gasinterface.

Michalek and Vacik in PCT/CZ2005/000093 describe an IOL manufacturingmethod using a spin-casting method in open molds. Molds filled withmonomer mixture spin along their vertical axis while polymerizationproceeds. One of the optical surfaces is created as the imprint of asolid mold surface while the other is formed by the mold rotation. Theimprinted surface has the shape formed by rotation of the conic sectionalong the vertical axis (which may include hyperboloid shape). The othersurface is shaped as a meniscoid modified by the centrifugal force thatwill transfer some of the liquid precursor from the center toward theperiphery. In the case of the convex meniscus, the centrifugal forcewill flatten the center and create a steeper curvature in the periphery,i.e. increase the spherical aberration of the surface. In the case of aconvex meniscus, the centrifugal force will create a meniscus withsmaller central radius and modify the surface to approximate somethingbetween spheric and parabolic shape. In any case, the hyperbolicaberration cannot be achieved for either a convex or concave meniscoidsurface.

Sulc et al. in U.S. Pat. Nos. 4,994,083 and 4,955,903 discloses anintraocular lens with its anterior face protruding forward in order tobe in permanent contact with the iris that will center the lens. Bothposterior and anterior surfaces may have the shape obtained by rotationof a conical section around of the optical axis (sphere, parabola,hyperbole, ellipse). The iris-contacting part of the lens is a hydrogelwith very high water content (at least 70% and advantageously over 90%of water) that is inherently soft and deformable. Therefore, the opticalsurface deformed by the contact with iris cannot be exactly a conicsection surface, but a surface with a variable shape that will depend onthe pupil diameter, probably close to a sphere with a somewhat smallercentral radius. Namely, this situation is similar to the lens fromanother reference that achieves decrease in the central diameter bypressing a deformable gel-filled lens against a pupil-like aperture inan iris-like artificial element (Nun in U.S. Pat. No. 7,220,279). Nun'279 does not mention or imply use of hyperboloid optical surfaces.Cummings in US Pat. Publ. Nos. 2007/0129800 and 2008/0269887 discloses ahydraulic accommodating IOL in which a liquid is forced into theinternal IOL chamber by action of cilliary apparatus causing thus changeof the optical surface and accommodation.

Hong et al. in U.S. Pat. No. 7,350,916 and US Pat. Publ. No.2006/0244904 disclose an aspheric intraocular lens with at least oneoptical surface having a spherical aberration in order to compensate thepositive spherical aberration of the cornea. The negative sphericalaberration is achieved by hyperbolic shape of the optical surface.

Hong et al. in US Pat. Publ. No. 2006/0227286 discloses optimal IOLshape factors for human eyes and defines the optimum lens by a certainrange of “shape factors” from −0.5 to +4 (the shape factor being definedby Hong as the ratio of sum of anterior and posterior curvatures totheir difference), and at least one of the optical surfaces isadvantageously aspherical with conic constant between −76 and −27.

Hong et al. in U.S. Pat. No. 7,350,916 describes an IOL with at leastone of the optical surfaces having a negative spherical aberration in arange of about −0.202 microns to about −0.190 microns across the powerrange.

SUMMARY OF THE INVENTION

In at least one aspect, the present invention provides an artificiallens implantable into the posterior chamber of human eye for replacementof the natural crystalline lens, the lens (referring to the FIG. 3)having a main optical axis 1A; the central optical part 2 and theperipheral supporting part 3; the overall shape of the implant beingdefined by its anterior surface 4, posterior surface 5 and thetransition surface 6 between the upper boundaries of the anterior andposterior surfaces 7A and 7B of the implant; having the central anterioroptical surface 8A with boundary 9A and anterior apex 10A; the centralposterior optical surface 8B with boundary 9B and posterior apex 10B;and anterior peripheral supporting surface 11A and posterior peripheralsupporting surface 11B.

Artificial lens implantable into the posterior chamber of human eye forreplacement of the natural crystalline lens that simulates as closely aspracticable the shape, size, optical properties and material propertiesof an NCL while respecting the need for surgical implantation through asmall incision.

The artificial lens according to at least one embodiment of theinvention has at least the posterior surface approximating the shape andsize of the posterior surface of the natural lens in order to achievesubstantially complete contact with the posterior capsule of the eye. Asdefined in this context, the term “substantially” can mean that eitherat least about 90% of the posterior BAIOL surface is in contact with theposterior capsule, or at least about 75% of the posterior capsule (thathas diameter larger that the lens) is in contact with the posteriorsurface of the lens. At least the part of the artificial lens accordingto the invention that is contacting the posterior capsule is made from atransparent flexible hydrogel material approximating the optical,hydrophilic and electrochemical character of tissue forming the naturallens. The anterior side is designed to avoid a permanent contact withiris.

In at least one embodiment, the anterior surface is shaped to avoid apermanent contact with the iris with the anterior peripheral supportingsurface 11A being concave.

In at least one embodiment, the artificial lens according to theinvention has at least the major parts of its anterior and posteriorsurfaces, including both optical surfaces, defined by rotation of one ormore conic sections along the optical axis and formed by solidificationof a liquid polymer precursor in contact with a solid wall of a mold,preferably a hydrophobic plastic mold.

One aspect of the invention is directed to a hydrogel comprisingUV-absorbing dopant moieties and activator moieties that are negativelycharged at physiological pH, where exposure of the fully hydratedhydrogel to electromagnetic radiation results in two-photon absorptionwhich causes one or more structural changes in the hydrogel and a changein the refractive index. In one embodiment the hydrogel is a covalentlycrosslinked hydrogel. In one embodiment, the structural change isachieved without substantially changing the volume of the treatedhydrogel segment after it gets into equilibrium with the liquid mediumaround the lens. One convenient method to determine the volume changecan consist in the following procedure: several equal zones in thehydrogel (e.g., 50×50 microns) are treated with various laser settingsto achieve different phase shift in different zones. After allowing timesufficient to achieve equilibrium, the linear dimension of the treatedzones does not change by more than 20% from the original zonedimensions, the dimension change being less than 10% for mostconditions. Since the depth of the treated zone cannot be readily ordirectly measured, it is assumed that the change of the volume isisotropic and the change of treated volume is due to the same relativeexpansion or contraction for all dimensions. One embodiment of thehydrogel comprises a polymer comprising monomer units of (meth)acrylicacid derivatives and/or (meth)acrylic acid. In one embodiment of thehydrogel the dopant moieties and activator moieties are pendant groupson a polyacrylate or polymethacrylate polymer. In one embodiment thedopant moiety is a UV-absorbing compound that does not strongly absorblight of about 400 nm wavelength. In one embodiment the dopant moiety isa compound selected from the group consisting of rhodamines,benzophenones, coumarins, fluoresceins, benzotriazoles, and derivativesthereof. In one embodiment, the UV absorbent moiety contains a carbonylgroup conjugated with an aromatic system, a phenolic hydroxyl groupconjugated with an aromatic system, or advantageously both carbonyl andphenolic hydroxyl groups conjugated with an aromatic system. In oneembodiment the activator moiety is a compound comprising a carboxylategroup, sulfonate group, sulfate group, phenolate group or phosphategroup. In one embodiment of the hydrogel the one or more structuralchanges comprise partial depolymerization of the hydrogel. In oneembodiment the partial depolymerization forms an aqueous-filled void inthe hydrogel. In one embodiment of the hydrogel the change in therefractive index is a negative change. In one embodiment the depth ofthe partial depolymerization of the hydrogel depends on the cumulativeenergy absorbed in a given location of the hydrogel. In one embodimentthe hydrogel comprises a polymer comprising monomer units selected fromthe group consisting of acrylic acid derivatives, methacrylic acidderivatives, acrylic acid, methacrylic acid, and mixtures of two or morethereof.

Another aspect of the invention is directed to an ophthalmic implantcomprising the above hydrogel.

Yet another aspect of the invention is directed to an in situ-adjustablehydrogel ophthalmic implant comprising an acrylate or methacrylatecopolymer hydrogel where the copolymer comprises at least fourco-monomers: a) an acrylate or methacrylate ester containing at leastone pendant hydroxyl group; b) a polyol acrylate ester or polyolmethacrylate ester or amide with at least 2 acrylate or methacrylategroups per polyol ester or amide; c) a derivative of acrylic ormethacrylic acid having at least one pendant carboxyl group; and d) avinyl, acrylic or methacrylic monomer having a pendant UV-absorbinggroup; where the refractive properties of the implant are adjusted by acontrolled absorption of targeted electromagnetic radiation by thehydrogel resulting in a change of refractive index in selected locationsof the implant. In one embodiment the ester of component a) carries atleast one pendant hydroxyl group on the alcohol portion of the ester. Inone embodiment the copolymer is covalently crosslinked. In oneembodiment the implant has anterior and posterior refractive surfacesforming a lens with positive or negative refractive power. In oneembodiment the lens is an intrastromal lens. In another embodiment thelens is an anterior chamber lens. In yet another embodiment the lens isa phakic lens for location between the iris and the natural crystallinelens. In a further embodiment the lens is a posterior chamber lens forat least partial replacement of the natural crystalline lens. In oneembodiment of the hydrogel ophthalmic implant at least one of therefractive surfaces is an aspheric surface with negative sphericalaberration. In another embodiment of the hydrogel ophthalmic implant themonomer containing the pendant carboxyl group is a neutralized orpartially neutralized methacrylic acid. In one embodiment the monomercontaining the pendant carboxyl group is present in a concentrationbetween 0.1 molar % and 5 molar % based on all monomer units of theco-polymer. In another embodiment the monomer containing the pendantcarboxyl group is present in a concentration between 0.5 molar % and 2molar % based on all monomer units of the co-polymer. In one embodimentthe monomer containing the pendant UV-absorbing group is present in aconcentration between 0.1 molar % and 5 molar % based on all monomerunits of the co-polymer. In another embodiment the monomer containingthe pendant UV-absorbing group is present in a concentration between 0.2molar % and 2.5 molar % based on all monomer units of the co-polymer. Inone embodiment of the hydrogel ophthalmic implant the pendantUV-absorbing group contains a carbonyl group conjugated to an aromaticgroup; in one embodiment the monomer containing the pendant UV-absorbinggroup is present in a concentration between 0.1 molar % and 5 molar %based on all monomer units of the co-polymer. In one embodiment of thehydrogel ophthalmic implant at least one of the UV-absorbing pendantgroups is selected from the group consisting of derivatives ofbenzophenone, derivatives of benzotriazole, derivatives of coumarin andderivatives of fluorescein. In one embodiment the pendant carboxylgroups and pendant UV-absorbing groups are present in a molar ratiobetween about 0.25 and 5; in another embodiment the pendant UV-absorbinggroups are present in a molar ratio between about 0.5 and about 3.5. Inone embodiment of the hydrogel ophthalmic implant the copolymer containsat least two different comonomers containing different UV-absorbinggroups; in one embodiment at least one of the UV-absorbing groups is abenzophenone.

In one embodiment of the ophthalmic lens the pendant carboxyl group isionized, and the molar ratio of ionized pendant carboxyl groups toUV-absorbing pendant groups is from about 0.5 to about 3.5. In anotherembodiment of the ophthalmic lens at least the major portion of thepolymer of the hydrogel is a derivative of methacrylic acid; in oneembodiment at least the major portion of the methacrylic acid derivativeis a hydrophilic derivative of methacrylic acid. In one embodiment thehydrophilic methacrylic acid derivative is a glycol ester of methacrylicacid. In one embodiment of the ophthalmic lens the covalentlycrosslinked hydrogel contains more than 30% by weight of liquid underequilibrium physiological conditions. In one embodiment the covalentlycrosslinked hydrogel contains less than 55% by weight of liquid underequilibrium physiological conditions. In another embodiment thecovalently crosslinked hydrogel contains between 35% and 47.5% by weightof liquid under equilibrium physiological conditions. In one embodimentof the ophthalmic lens at least the posterior optical surfacecontributes refraction with negative spherical aberration. In anotherembodiment the posterior optical surface contributes refraction withnegative spherical aberration between −0.1 microns and −2 microns. In afurther embodiment the negative spherical aberration is between −0.5microns and −1.5 microns. Alternatively the negative sphericalaberration is between −0.75 microns and −1.25 microns. One embodiment ofophthalmic lens comprises both UV-absorbing pendant groups containingcarbonyl groups conjugated to an aromatic system as well as UV-absorbinggroups containing a benzotriazole structure. In one embodiment theUV-absorbing pendant groups containing carbonyl groups conjugated to anaromatic system and the UV-absorbing groups containing a benzotriazolestructure are located in separate layers of the ophthalmic lens. In oneembodiment the ophthalmic lens is implanted into a cornea. In anotherembodiment the lens is implanted into the anterior chamber of an eyebetween the cornea and iris. In yet another embodiment the lens is aphakic lens implanted between the iris and the natural crystalline lens.In still another embodiment the lens is implanted into the posteriorchamber of eye, at least partially replacing the natural crystallinelens.

Another aspect of the invention is directed to a method of adjusting therefractive properties of a fully hydrated hydrogel of the invention,where the method comprises the step of focused irradiation of thehydrogel with electromagnetic radiation such that two-photon absorptionoccurs, and the polymer component of the hydrogel undergoes partialdepolymerization and/or decomposition, with effective removal of part ofthe polymer scaffold to create a void.

An additional aspect of the invention is directed to a method of in situadjusting the optical parameters of a hydrogel ophthalmic implant, themethod comprising the steps of: a) providing an eye containing ahydrogel ophthalmic implant according to claim 13; and b) irradiating aportion of the hydrogel ophthalmic implant with electromagneticradiation using a femtosecond laser, whereby part of the copolymer ofthe hydrogel is depolymerized and/or ablated; where the opticalparameters of the implant are adjusted. In one embodiment of the methodthe optical parameters include the refractive index. In one embodimentof the method the irradiation produces elongated cavities or voxelsinside the hydrogel ophthalmic implant. In some embodiments the voxeldepth is up to 20-30 microns or more. In one embodiment increasing thedepth of the voxels increases the phase shift while the refractive indexremains approximately constant; in one embodiment the refractive indexis 1.3335. In some embodiments of the method the phase shift is up to 3wavelengths of green light. In other embodiments of the method thedepolymerized matter ablated by irradiation comprises soluble, easilydiffusible compounds of low toxicity. Since the procedure releases onlyvery low concentrations of depolymerized compounds into the intraocularspace, the low toxicity of the method is also evident. In one embodimentof the method the modified optical properties are provided by forming inthe hydrogel a pattern of elongated voxels of varying depth whilekeeping the modified refractive index approximately constant. In oneembodiment of the method the phase shift is controlled by varying voxeldepth rather than by varying their refractive index.

BRIEF DESCRIPTION OF THE DRAWINGS

The accompanying drawings, which are incorporated herein and constitutepart of this specification, illustrate the presently preferredembodiments of the invention, and, together with the general descriptiongiven above and the detailed description given below, serve to explainthe features of the invention. In the drawings:

FIG. 1 illustrates the internal arrangement of the eye with mainstructures including the cornea, sclera, iris, NCL, vitreous body,retina and the suspensory apparatus of the lens (capsule, zonules andcilliary muscle)

FIG. 2 illustrates distribution of refractive power in a lens with onehyperbolic surface.

FIG. 3A is a cross-sectional view of a bioanalogic intraocular lensaccording to an exemplary embodiment of the invention.

FIG. 3B is a top view of the lens of FIG. 3A.

FIG. 4A is a top view of another exemplary embodiment of a lens with acircular optical part and elliptical support part.

FIG. 4B is a top view of another exemplary embodiment of a lens with acircular support part truncated by a single straight cut.

FIG. 4C is a top view of another exemplary embodiment of a lens with acircular support part truncated by two symmetric crescent cuts.

FIG. 4D is a top view of another exemplary embodiment of a lens with acircular support part truncated by one straight and two crescent cuts.

FIG. 4E is a top view of another exemplary embodiment of a lens with acircular support part truncated by four symmetric crescent cuts.

FIG. 4F is a top view of another exemplary embodiment of a lens with acircular support part truncated by two straight parallel cuts and thecylindrical lens with cylinder axis 1B in the angle α with regard to thecuts direction.

FIGS. 5A, 5B and 5C illustrate top views of exemplary lenses with theoptical surfaces divided into two or more optical zones.

FIGS. 6A, 6B and 6C are cross-sectional views of alterative lens inaccordance with the invention composed from two or more materials.

FIGS. 7A, 7B and 7C are expanded views illustrating alternative profilesof the supporting peripheral part of the exemplary lenses.

FIG. 8 illustrates the schematic arrangement of the mold for productionof a lens in accordance with an exemplary embodiment of the invention.

FIG. 9A shows a comparison of the Raman spectra of a representativehydrogel of the invention before and after two photon absorption (TPA)using a femtosecond laser.

FIGS. 9B and 9C show graphs of relevant parameters of the Ramanspectrum.

DETAILED DESCRIPTION OF THE INVENTION

For the purposes of the present application, the terms “(meth)acrylic”and “(meth)acrylate” denotes either acrylic/acrylate ormethacrylic/methacrylate moieties. In some embodiments the polymerscomprise acrylic/acrylate moieties. In other embodiments the polymerscomprise methacrylic/methacrylate moieties. In still other embodimentsthe polymers comprise both acrylic/acrylate and methacrylic/methacrylatemoieties. In a preferred embodiment the polymers comprisemethacrylic/methacrylate moieties.

Also, for the purposes of the present application, the term “ablation”refers to removal of a section of the polymeric structural support of ahydrogel, preferably by decomposition of the polymer to diffusible,low-molecular fragments. The term “depolymerization” constitutes aspecial type of ablation where the fragments are monomers.

There are numerous types of implantable ophthalmic lenses used invarious locations in the eye, from corneal stroma through anteriorchamber to posterior chamber. One of the problems with implantablelenses is the complicated selection of the correct refractive properties(so called biometry) and difficulty to replace them or correct them ifthe biometry turned out to be wrong or if the optical requirements ofeye change over time. This inspires the effort of the industry todevelop implantable lenses whose optical parameters could be adjustednon-invasively and post-operatively. Change of the optical properties ofan artificial lens, either preoperatively or post-operatively, can beachieved by changing refractive index of the lens material.

The change of the refractive index by action of Two Photon Absorption(TPA) or Multi-Photon Absorption (MPA) for implantable lenses isdescribed in prior art as positive for hydrophilic polymers and negativefor hydrophobic polymers. This is consistent with assumed mechanism ofincreased crosslinking density and consequent decrease of water contentin hydrophilic acrylates and hydrogels, and conversely, with increasedhydrophilicity in the hydrophobic acrylates. The assumed cause of thesechanges is a local increase of temperature.

Change of the optical properties of the natural (e.g., human cornea) oran artificial lens by changing refractive index of their material isalso described in numerous papers, patents and patent applications, suchas the following: Phillips, A. J., System and Method for Treatment ofHyperopia and Myopia, U.S. Pat. No. 6,102,906, Bille J. F.: System ForForming And Modifying Lenses And Lenses Formed Thereby, U.S. Pat. Nos.8,292,952; 8,920,690; 9,192,292; Sahler; Ruth et al.: “Hydrophilicityalteration system and method” U.S. Pat. Nos. 9,023,257; 9,186,242 and9,107,746; Sahler; Ruth et al.: “intraocular lens (IOL) fabricationsystem and method” US Patent Application Publication No. 20160074967;Smith, T. et al.: Optical Hydrogel Material With Photo sensitizer AndMethod For Modifying The Refractive Index, US Patent ApplicationPublication Nos. 20130268072; 20090287306 and U.S. Pat. No. 8,901,190;Knox, Wayne H. et al.: Optical Material And Method For Modifying TheRefractive Index, U.S. Pat. Nos. 8,932,352; 8,337,553; 7,789,910 B2;Knox, Wayne H. et al.: Optical Material And Method For Modifying TheRefractive Index, US Patent Applications Publication Nos. 20130138093;20130178934; 20100298933; 20080001320; 20090143858; 20090143858; Knox,Wayne H. et al.: Method For Modifying The Refractive Index Of An OpticalMaterial And Resulting Optical Vision Component, US Patent ApplicationPublication No. 20120310340, Knox, Wayne H. et al.: Method For ModifyingRefractive Index Of Ocular Tissues, U.S. Pat. Nos. 8,486,055; 8,512,320;8,617,147 and US Patent Application Publication Nos. 20110071509;20130226162 and 20140107632; Knox, Wayne H. et al.: Method For ModifyingRefractive Index Of Ocular Tissues And Applications Thereof, US PatentApplication Publication No. 20120310223, each of which is incorporatedherein by reference. None of these prior art references suggests anablation or depolymerization as the mechanism of the optical adjustmentin hydrogels.

Applicant's co-pending international application, PCT/IB2016/052487,Method and Device for Optimizing Vision Via Customization of SphericalAberration of Eye, filed on May 2, 2016, contains related disclosure.

The wavelength of the laser beam is usually in the range of nearinfrared radiation, about 800 nm to 1300 nm, or more typically in therange of visible and near infrared radiation from about 660 nm to about1100 nm. The use of higher wavelengths is usually preferred because ofsafety concerns. One method of changing the refractive index of the lensmaterial, by means of a femtosecond laser (“FSL procedure” for short),can, in principle, achieve many changes of refractive properties, suchas spherical refractive power, cylindrical refractive power, sphericalaberration, etc. In principle, the procedure can selectively change anyof the coefficients in the Zernike polynomial (“Zernike Coefficients”)repeatedly even in the implanted lens. This has been demonstrated byGustavo A. Gandara-Montano et al “Femtosecond laser writing of freeformgradient index microlenses in hydrogel-based contact lenses”, 1 Oct.2015|Vol. 5, No. 10|DOI:10.1364/OME.5.002257|OPTICAL MATERIALS EXPRESS2257. The authors noted a negative phase shift in the ETAFILCON® contactlens hydrogel when treated by femtosecond laser at 800 nm. However, noneof these systems is in clinical use because, among other reasons, thechange of the refractive index using an FSL procedure is rather smallfor the currently available ophthalmic lens materials.

The FSL procedure can be performed both on hydrophilic (where therefractive index RI is usually increased) and hydrophobic IOL materials(where the RI is usually decreased).

The FSL procedure can be more advantageously performed on hydrogelophthalmic lenses, particularly on implantable lenses of various types,since the refractive change even in the currently available hydrogels ishigher than in the hydrophobic materials. In addition, FSL procedure inhydrophobic acrylates could lead, at least in theory, to so-called“glistening” or other problems related to formation of hydrophilic“osmotic cells” within a hydrophobic material. Therefore, varioushydrogels were tested as substrates for micromachining by femtosecondlasers. The sensitivity of hydrogels can be increased by various“dopants” designed to respond to various wavelengths of electromagneticradiation. A number of dopants have been described so far in the patentand scientific literature. All presently known dopants are singlecompounds capable of UV absorption, but no known dopants are capable ofsingle-photon absorption at the wavelength used for the FSL procedure.

We have found that the effect of dopants for TPA and MPA in the acrylicand methacrylic hydrogels can be improved, modified and strengthened byadding certain TPA activators, e.g. co-monomers containing negativelycharged pendant groups, particularly organic carboxylic acid salts. Boththe dopant and its activator can be advantageously covalently bound topolymer chains, more advantageously to the chains forming the polymernetwork of the hydrogel. Both the dopant and its activator may be boundto the same polymer chain or to different polymer chains.

More particularly, the inventive hydrogels contain a combination dopantin the form of a minor part of a pendant UV-absorbing structure thatdoes not significantly absorb visible light, with a minor part of theactivator in the form of pendant groups comprising an ionized salt of anacid, with a major part of methacrylate neutral hydrophilic derivative,particularly glycol or glycerol esters of methacrylic acid and a minorpart of a polyol with at least two of the hydroxyl groups esterified byacrylic or methacrylic acid. These hydrogels are materials particularlysuitable for adjustment of refractive index by absorption of femtosecondpulses of visible or near infrared radiation (femtosecond laser (FSL)treatment). The absorption of electromagnetic energy causes controlleddegradation and depolymerization of the polymeric component of suchhydrogels thereby forming domains with lower refractive index. Thanks tothe activator, such domains can be formed even if the hydrogel isirradiated by a femtosecond laser with pulses of relatively low energyand at a very high “writing speed” or “scanning velocity”.

Disclosed herein are bioanalogic covalently crosslinked acrylic andmethacrylic hydrogels containing negatively charged groups, particularlycarboxylate groups, and containing also monomers with pendantUV-absorbent groups, such as, for example, methacryloyloxybenzophenone(MOBP). Also disclosed herein is a method to depolymerize such polymersby absorption of electromagnetic radiation in order to adjust opticalproperties of the implantable hydrogel using femtosecond laser todepolymerize parts of the hydrogel and to form internal cavities thatwould form new refractive members, such as toric lenses. Implicitly butcertainly, such newly formed cavities in the hydrogel matrix would benecessarily filled with water or aqueous fluid (either alone or alsocontaining the residual of the degraded polymer component of thehydrogel) and would have, therefore, different (specifically lower)refractive index than the parent hydrogel material. Also disclosed isthe possibility to perform this lens adjustment in situ on the implantsince the products of such a decomposition are water-soluble compoundsof low toxicity and capable of a slow diffusion through the hydrogel.

We have confirmed that the UV-absorbent group used in the hydrogel actsas a “dopant” increasing the depolymerization rate, while the negativelycharged group acts as an “activator” for the dopant, increasing thedopant's efficacy still further. In addition, the activator acting as aquencher protects the material from undesirable “charring” or “burning”that may happen over a certain amount of absorbed energy per time. Thisallows one to lower the overall amount of energy needed to achieve acertain refractive change, and therefore to use safely the moreenergetic visible light rather than less energetic near infrared (NIR)radiation.

Without wishing to limit the scope of the invention by any particularhypothesis or theory, the proposed explanation of the phenomena of thehydrogel depolymerization by electromagnetic radiation is as follows.Focused femtosecond lasers are known to facilitate two-photon-absorption(TPA) (or even Multi-Photon Absorption (MPA) in their focal volume. Thevolume of the material affected by the TPA is usually called “volumepixel”, or “voxel”. The size of the voxel depends on various parameters,and generally increases with the absorbed energy. The smallest possiblevoxel size corresponds to the “focal volume”, or an ellipsoid withvolume roughly the cube of the wavelength of the laser light absorbed.The diameter of a voxel is much smaller than its depth thanks to theself-focusing of the laser beam in the case of TPA or MPA, and istypically around 500 nm. The voxel depth is much larger and grows withdeposited energy. The reported maximum voxel depth is about 6 microns incurrently known materials.

If the radiation of a certain wavelength is absorbed via TPA or MPA,then the dopant accumulates excitation energy equivalent to asingle-photon absorption of the wavelength of incident light divided bythe number of the absorbed photons. For example, laser light at 400 nmabsorbed via TPA by the dopant will correspond to 200 nm light absorbedvia SPA. Light of 200 nm wavelength is a hard UV light (UVC Band) thatis sufficiently energetic to cause a breakage of chemical bonds and adeep structural rearrangement. Of course, 3 photon absorption or 4photon absorption, although much less probable than TPA, wouldaccumulate even higher energy concentration. The excited state with highaccumulated energy is short-lived since the energy dissipates quicklyvia one of several possible pathways, the most usual being conversioninto thermal energy. The usual effect of femtosecond laser treatment oftissues or synthetic hydrogels involves local increase of temperaturethat may cause—depending on the amount of the energy absorbed—effectsstarting with conversion of the matter into a plasma (e.g., via laserablation) through heating sufficient for charring, to more subtleadditional crosslinking via various mechanisms such asre-esterification, disproportionation or dehydration with formation ofether links.

However, the presence of the activator (pendant negatively chargedgroups, particularly carboxylate groups), changes the mechanism of thisprocess. We believe that this is caused by certain cooperation of theactivator groups with dopants. In this cooperation, activation ofdopants further facilitates TPA absorption and increases its effect. Theactivator, such as the carboxylic acid group of methacrylic acid or asalt thereof, is apparently in an interaction with the dopant becauseits presence causes a change of the dopant's UV/Vis spectrum (namely, asubtle shift from the UV toward the visible region). This may increasethe dopant's “TPA cross-section” and increase the absorption efficacy ofthe dopant.

However, the main role of the activator may be in channeling theabsorbed energy of the two (or more for MPA) absorbed photons from thedopant to the main polymer chain to cause breakage of the covalent bondin the main chain. It is understood from the general character of thisprocess, and from the known depolymerization kinetics of methacrylatepolymers, that this cleavage is homopolar and produces free radicalsthat start the depolymerization process in the vicinity of thenegatively charged activator (e.g., carboxylate). This free radicalmechanism may vastly increase the quantum yield of the photodegradationand helps to explain why even a relatively low concentration (on molarbasis) of the dopant can cause profound changes in the local compositionand structure, and consequently a substantial change in the localrefractive index.

Furthermore, the activator groups help to divert the absorbed energyfrom the dopant to quench the excited state and “consume” this energyfor the breakage of covalent bonds. This “energy conduit” function ofthe activator prevents excessive heat accumulation in the vicinity ofthe dopant moiety, helping thus to “recycle” the dopant molecule andpreserve it for the future TPA cycle. It also helps to protect thepolymer structure from burning and charring and helps to increase theefficacy of the whole absorption process.

In addition, this proposed mechanism can explain the fact that in theabsence of the activator, TPA in hydrogels with only the dopant proceedsvia a different mechanism and achieves the opposite result: while inpresence of the activator the refractive index decreases, in its absencethe refractive index increases. One possible reason for the increasedrefractive index seems to be an additional crosslinking leading todecreased water content and, consequently, increased refractive indexsince water has the lowest refractive index of all hydrogelconstituents.

Another consequence of the stipulated mechanism involving the activatoris the reduced (or even eliminated) change of the volume due the TPAprocess. Namely, in the case of the additional crosslinking the watercontent is reduced while the amount of the polymer components staysapproximately the same. Therefore, the mass and the volume of thetreated hydrogel volume has to be reduced, with various adverseconsequences on the product (generation of internal stress, change ofgeometry and change of mechanical properties, for instance). Conversely,if the polymer becomes more hydrophilic and attracts more water, thetreated volume of the polymer has to expand (again, with adverse effecton the geometry, optics and stresses within the material).

The mechanism stipulated in the present invention involves exchange of acertain part of the polymer mass for water, with a low net volumechange, if any. The depolymerization of the hydrogels covered by thepresent invention yields low toxicity monomers and/or their fragments,primarily 2-hydroxyethyl methacrylate, methacrylic acid and ethyleneglycol. All these decomposition products are well soluble in water andcapable of diffusing through the surrounding intact hydrogel network, tobe dissipated from the implant over some time in very lowconcentrations.

The proposed mechanism may explain very large phase shift achievable inhydrogels according to this invention. The phase shift is determined bythe change in the refractive index and the length of light path throughmaterial with changed refractive index, in other words, the voxel depth.Refractive index in hydrogels cannot be practically lower thanrefractive index of water, or about 1.3335 since the lowest conceivableindex can be achieved in a cavity filled with aqueous liquid. The cavityin a hydrophilic polymer matrix cannot be filled with a gas for certainbasic thermodynamic reasons, and there is no known or readilyconceivable mechanism of creating a hydrophobic cavity within thehydrogel.

Therefore, large phase shifts (more than one wavelength) will requireformation of voxels with large depth. It is usually assumed that thevoxel depths can reach at most about 5 to 6 microns. However, if therefractive index in the focal volume decreases, then the elongated voxelwould act as a light-guide that channels additional light pulses to beabsorbed in the voxel bottom. Consequently, the depth of the voxel cangradually increase with increasing number of pulses, and so does thenthe phase shift even though the refractive index remains approximatelyconstant and equal to or higher than 1.3335. This mechanism is thendifferent from mechanisms described to date where voxel size remainsapproximately constant but refractive index change increases withincreased absorbed energy (achieved e.g., number of pulses). It isbelieved that the presently proposed mechanism may provide a voxel depthlarger than 10 microns, and phase shifts corresponding to voxel depth ashigh as 25 or 30 microns have been observed. This is mentioned not toset a limit on the invention, but demonstrate the fundamental differenceof the invention from prior art. One of the practical consequences ofthis difference is the following: materials used in the prior art allowfor creation of refractive or diffractive structures by forming arefractive index gradient (GRIN) while hydrogels according to thepresent invention allow achieving similar refractive or diffractiveeffects by forming a pattern of varying voxel depth while keeping themodified refractive index approximately constant. In addition, it isbelieved that this refractive index value first approaches andultimately approximates the refractive index of isotonic salinesolution.

As a consequence of this novel mechanism, the newly created refractiveor diffractive structures form a system of parallel waveguides (i.e.,elongated voxels with refractive index lower than the surroundinghydrogel) of different length, whereby the phase shift is controlled bythe varying voxel depth rather than their varying refractive index.

An additional feature of the invention is creation of a gradient ofrefractive index in the vicinity of individual voxels. The monomers andother fragments released by the depolymerization inside the voxelmigrate radially by diffusion through the surrounding hydrogel. When thetemperature decreases below the ceiling temperature (around 200° C. inthe case of methacrylates), at least part of the released monomersand/or fragments may re-polymerize and create a denser network structurewith a higher refractive index than the parent hydrogel. This mechanismhas two beneficial consequences: first, it reduces the amount ofcompounds that diffuse outside of the implant to be metabolized, andsecond, the gradient of refractive index thus formed improves thelight-guiding properties of the voxel.

Thus, monomers released by depolymerization may not all diffuse out ofthe hydrogel, but may partly re-polymerize in the voxel vicinity. Inthat sense the present method is somewhat similar to FS laser ablationin that that both remove some polymer mass and retain water instead,although the term “ablation” frequently denotes a process ofdecomposition where the polymer and water are converted into a plasma(gas). The presently disclosed process is much more gentle so thatpolymer decomposition does not create an exploding bubble of gas.

Although depolymerization of the hydrogel copolymer is believed to bethe primary mechanism involved in the presently disclosed process, itmay be supplemented by other decomposition reactions that form small,water-soluble fragments, such as hydrolysis of the (meth)acrylatependant groups, or oxidation.

Dopant concentration in hydrogels according to the invention varies fromabout 0.05%-mol to about 5%-mol, advantageously from about 0.1%-mol toabout 2.5%-mol. Alternatively the different optimum dopant concentrationcan be used for various dopants, e,g, 0.25 to 0.55 molar % forbenzophenone derivatives, or 0.1 to 0.2 molar % for benzotriazolederivatives. The preferred dopants are UV absorbers with low absorbancefor visible light, i.e. above a wavelength of about 390 nm. Examples ofsuitable dopants are vinyl, acrylate or methacrylate derivatives ofbenzophenone, benzotriazole and coumarin, although those skilled in theart can certainly identify other suitable UV absorbers that work asdopants in the sense of the invention.

The activating groups are present in concentration from about 0.25%-molto about 5%-mol, advantageously from about 0.75 to about 3.5%-mol. Insome embodiments the preferred molar concentration of methacrylic acidis between 1 and 1.25% molar. The preferred activator groups arederivatives of acrylic or methacrylic acid containing a pendant acidicgroup including carboxylate group, sulphate group, sulfonate group andphosphate group. Such acidic groups are preferably neutralized bysuitable organic or inorganic cations. The activator to dopant groupmolar ratio should be about 0.75 to 10, advantageously from about 1 toabout 5. Alternatively the different optimum molar ratioactivator/dopant are different for different dopant. For instance, theoptimum ratio for benzophenone derivatives is from about 2 and 4 and forbenzotriazol derivatives between about 5 and about 7.

The preferred composition of a hydrogel according to the inventioncomprises a major fraction of a methacrylate monomer with pendantneutral hydrophilic groups. By “major fraction” is meant at least50%-mol of all monomer units in the hydrogel. In some embodiments themolar fraction of the hydrophilic methacrylate monomer is between90%-mol and 99.5%-mol, and in many cases between 97.5%-mol and 99%-mol.Such hydrophilic methacrylate monomers can be esters of polyol aliphaticcompounds, such as glycols, glycol ethers, glycerol and sugars. The mostfamiliar in this group is 2-hydroxyethylmethacrylate (2-HEMA).Alternative monomers to the above mentioned esters comprise amides ofmethacrylic acid, such as methacrylamide, N-isopropyl methacrylamide orN-(2-hydroxyethyl) methacrylamide. This major fraction of the hydrogelpolymer may also comprise a mixture of such hydrophilic monomers.Alternatively, a minor portion of the methacrylate monomers (but notmore than 25%-mol) can be replaced by analogous acrylic acidderivatives. In some embodiments 0.5%-mol to 5%-mol can be replaced byacrylate monomers.

A minor part of the monomer units is formed by the above mentionedactivator monomers with a pendant negatively charged group, and stillanother minor part is formed by the above mentioned dopant monomers.

The polymer is advantageously covalently crosslinked. The crosslinkingcan be achieved by any of the methods known to those skilled in the art,such as radiation crosslinking, crosslinking by formation of ether linksbetween pendant OH groups, etc. The preferred crosslinking method iscopolymerization with a minor fraction of methacrylate or acrylatecrosslinking diesters or triesters of polyols, such as, for example,triethylenglycol dimethacrylate.

Refractive index of the hydrogel according to the invention is betweenabout 1.38 and about 1.48, preferably between about 1.40 and about 1.45.In some embodiments the RI is about 1.40, or about 1.41, or about 1.42,or about 1.43, or about 1.44, or about 1.45.

The preferred hydrogel according to the invention contains between about25%-wt and about 85%-wt liquid in equilibrium with live intraocularenvironment. For intraocular lenses supplementing or replacing thenatural crystalline lens, the more advantageous equilibrium liquidconcentration is between 35%-wt and 50%-wt, and more particularlybetween 40% and 47%-wt. In some embodiments the equilibrium watercontent is 41%-wt±0.75%, or 42.5±1%-wt, or 44.5±1% by weight. It isunderstood by those skilled in the art that equilibrium liquid contentin hydrogels is subject to many variables, such as body temperature,composition of body fluids or pressure exerted by surrounding tissues orbody structures. Also, measurement of liquid content in small hydrogelimplants may be burdened by a certain measurement error, so that thesevalues are illustrative. Since the negatively charged activator groupstend to increase equilibrium liquid content and thus to decrease therefractive index of the hydrogel, the higher concentration of theactivator may be compensated with addition of hydrophobic acrylates ormethacrylates, such as methylmethacrylate, ethylmethacrylate,benzylmethacrylate, isobornylmethacrylate or ethoxyethyl methacrylatethat will reduce the water content and increase refractive index into adesirable range for the ophthalmic implant design. Typical concentrationof such added methacrylates or acrylates is up to about 40% molar. Insome embodiments the concentration of added hydrophobic monomers isbetween about 5%-molar and 25%-molar.

The dopant group and activating group may be located on the same polymerchain, or on different polymer chains that are in an intimate contacte.g. in a polymer blend, or being on two different segments of polymernetwork. One or other may be on a graft, or one can be on a graft andthe other on the basic chain. The dopant group and activating group maybe on a single molecule in a suitable steric relationship allowing theirmutual interaction. Such a “dopant-activator complex” may be thencovalently bound to a polymer chain, or admixed into the polymer, orbecome part of the main polymer chain.

Negatively charged activator groups have some additional advantages forintraocular implants, such as improved biocompatibility, resistance toadsorption of proteins, resistance to formation of biofilms, resistanceto calcification, and resistance to adhesion and spreading of cells(which translates into resistance to sclerotization of posterior capsuleand resistance to posterior capsule opacification).

The hydrogel implant according to the invention can be placed in variouslocations within the eye along the optical path. It may have the form ofa “blank” in which a refractive or diffractive lens is created, or itmay have the form of a refractive or diffractive lens which opticalproperties are modified by the refractive index change within selectedlocations of the implant.

Intraocular lenses may replace partly or fully the natural crystallinelens. Intraocular hydrogel lenses have some advantages over other IOLtypes, and various types and related manufacturing methods aredescribed, e.g., in the following patents and patent applications: Stoy,V. et al.: Bioanalogic Intraocular Lens, International PatentApplication WO2014111769; Wichterle, O.: Method Of Molding AnIntraocular Lens, U.S. Pat. No. 4,846,832; Wichterle, O.: Soft AndElastic Intracameral Lens And Method For Manufacturing Thereof, U.S.Pat. No. 4,846,832; Stoy, V.: Implantable Ophthalmic Lens, A Method OfManufacturing Same And A Mold For Carrying Out Said Method, U.S. Pat.No. 5,674,283; Sulc, J., et al.: Soft Intracameral Lens, U.S. Pat. Nos.4,994,083 and 4,955,903; and Michalek, J. et al.: Method OfManufacturing An Implantable Intraocular Planar/Convex, Biconvex,Planar/Concave Or Convex/Concave Lens And A Lens Made Using This Method,U.S. Pat. No. 8,409,481; each of which is incorporated herein byreference.

Another type of implantable ophthalmic lens is so called “ImplantableContact Lens” (or ICL). ICL is a phakic lens placed between iris and thenatural crystalline lens. It is described in several patents, e.g.,Fedorov, et al. Intraocular lens for correcting moderate to severehypermetropia, U.S. Pat. No. 5,766,245; Feingold V., Intraocular contactlens and method of implantation, UP 5,913,898; Intraocular refractivecorrection lens, U.S. Pat. No. 6,106,553; each of which is incorporatedherein by reference.

These implantable lenses are made from hydrogels that incorporate abiological component, usually collagen. These so called “Collamers” aredescribed in various patents, such as Feingold, et al., Biocompatible,optically transparent, ultraviolet light absorbing, polymeric materialbased upon collagen and method of making, U.S. Pat. No. 5,910,537;Feingold, et al. Biocompatible optically transparent polymeric materialbased upon collagen and method of making, U.S. Pat. Nos. 5,654,349,5,654,388 and 5,661,218; Fedorov, et al., Biocompatible polymericmaterials, methods of preparing such materials and uses thereof, U.S.Pat. No. 5,993,796; Method of preparing a biological material for use inophthalmology; each of which is incorporated herein by reference.

Another approach to refractive correction is intrastromal orintracorneal implants or inlays, described e.g. by Miller in Asphericalcorneal implant, U.S. Pat. No. 7,776,086; Lang, Alan in Design of InlaysWith Intrinsic Diopter Power, US 2007/0255401; Dishler; Jon et al. inSmall Diameter Inlays, US 2007/0203577; Lang, Alan et al. inIntracorneal Inlays, US 2007/0129797; and Dishler, et al. in Method ofusing small diameter intracorneal inlays to treat visual impairment,U.S. Pat. No. 8,057,541; each of which is incorporated herein byreference.

Some intrastromal implants are designed for post-operative adjustment ofoptical power by laser, as described by Peyman, Gholam A. inIntrastromal corneal modification via laser, US 2001/0027314; inAdjustable ablatable inlay, US 2002/0138069 and 2002/0138070; Ablatableintracorneal inlay with predetermined refractive properties in US2003/0093066; and in Bifocal implant and method for altering therefractive properties of the eye, US 2005/0222679; each of which isincorporated herein by reference.

There are also intraocular implants using two lenses in tandem in orderto achieve improved accommodation capability, modularity of the designor reduced size of implantation incision when the implant is in situassembled from individual parts.

All such ophthalmic lens types can be manufactured from the hydrogelsaccording to the present disclosure and then modified by using afemtosecond laser treatment.

Again, without wishing to be limited by any particular theory, thepartial depolymerization of the presently disclosed method appears to bethe mechanism of the profound negative phase shift observed in theinventive hydrogels when exposed to focused femtosecond lasers (FSLs).As noted above, this process appears to generate a system of parallellongitudinal “voxels” that get more elongated with increasing number ofabsorbed FSP pulses. The voxel length controls the phase shiftachievable for a given refractive index, that cannot decrease below theRI value for water, which is 1.3335. Further, it is actually the phaseshift, not the refractive index per se, that is affected in thepresently disclosed hydrogels and methods. This mechanism is differentfrom anything disclosed by the prior art (presumed crosslinking leadingto decrease in the water content that leads to the refractive indexincrease and, therefore, a corresponding positive phase shift).

A further difference from the prior art is that the presently disclosedhydrogels and methods replace part of the hydrogel-forming copolymerwith water or aqueous fluid, rather than decreasing the water content ofthe hydrogel by modifying the copolymer properties (such as by polymercrosslinking).

In the drawings, like numerals indicate like elements throughout.Certain terminology is used herein for convenience only and is not to betaken as a limitation on the present invention. The following describespreferred embodiments of the present invention. However, it should beunderstood, based on this disclosure, that the invention is not limitedby the preferred embodiments described herein.

The NCL has a very complicated structure that develops over time. One ofthe structural features is asphericity of posterior and anteriorsurfaces of the NCL 103. As established in recent years E. L. Markwellet al, MRI study of the change in crystalline lens shape withaccommodation and aging in humans, Journal of Vision (20110 11(3); 19,1-16; M. Dubbelman et al, Change in shape of the aging human crystallinelens with accommodation, Vision Research 45 (2005), 117-132; F. Manns etal, Radius of curvature and asphericity of the anterior and posteriorsurface of human cadaver crystalline lens, Experimental Eye Research 78(2004), 39-51; M. Dubbelman et al, The shape of the aging human lens:curvature, equivalent refractive index and the lens paradox, VisionResearch 41 (2001) 1867-1877, both anterior and posterior surfaces of ayoung human lens are hyperbolic and can be characterized by theequation:

Y−Yo=X{circumflex over ( )}2/{Ro*(1+1−h*(X/Ro){circumflex over( )}2){circumflex over ( )}0.5}  equation 1

where Y is the coordinate in the direction of the main optical axis 1A,X is the distance from the main optical axis 1A, Yo is the apex positionon the main optical axis 1A, Ro is the central radius of curvature and his the conic constant (or the shape parameter). The Eq. 1 describes anyconic section curve depending on the shape parameter h value: it is aparabola for h=0, a circle for h=1, hyperbole for h<0, prolate ellipsefor 0<h<1 and oblate ellipse for h>1.

It has been found that for a typical young human NCL, the anteriorsurface is more hyperbolic than the posterior surface, thathyperbolicity increases significantly with accommodation, and that thehuman lens grows with age and its hyperbolicity decreases so that an oldNCL may become approximately spherical.

The referenced studies mapped dimensions of typical NCL for selectedpopulation samples. According to these references, a typical human lensanterior central radius ranges from about 5 to 13 mm and the averageanterior conic parameter is about −4 (ranging from about −22 to +6). Theposterior central radius ranges from about 4 to 8 mm and the averageposterior conic parameter is about −3 (ranging from about −14 to +3).

The central thickness of a young, relaxed NCL ranges typically fromabout 3.2 mm to about 4.2 mm, increasing with age and/or with thenear-focus adjustment to a thickness from about 3.5 mm to about 5.4 mm.The posterior part depth of the NCL is typically the same as, or largerthan the anterior part depth. Therefore, the sagittal depth of theposterior lens surface is typically from about 1.75 mm to about 2.75 mmon equatorial diameter from about 8.4 mm to about 10 mm. This definesthe basic dimensions of the posterior capsule in its “natural” state.

Although the above references do not state any particular connectionbetween the geometry and optical properties of the NCL, we have found bymathematical modeling that the hyperbolic surfaces turn a lenspolyfocal, with the refractive power maximum at its center and graduallydecreasing toward the periphery. One direct consequence expected fromsuch a polyfocality is a large focal depth of the lens so that a nearobject can be projected on the retina even without any particular lensshape change. Another implication of the modeling is that the averagerefractive power of the lens increases with decreasing aperture.Therefore, it is concluded that the near focus can be improved by pupilconstriction (this so called “pupillary reflex” or “near myosis” thatcan be actually clinically observed at near focus). Another consequenceof the natural lens hyperbolicity is the capability of the human (andparticularly young) brain to naturally neuro-adapt to, and correctlyinterpret images formed by projection through a hyperbolic lens ontoretina.

This accommodative mechanism utilizing certain type of the polyfocalitydeserves further explanation as follows.

Lenses with at least one hyperbolic surface demonstrate a “hyperbolicaberration” that is opposite of the spherical aberration: rays incomingthrough the center are bent into a focal point that is closest to thelens, and the focal point becomes progressively further from the lensfor rays incoming in increasing distance from the lens center toward thelens periphery.

Therefore, the lens with hyperbolic surface is positively polyfocal: ithas the shortest focal distance (i.e., highest refractive power) at itscenter, and the focal distance increases (i.e., refractive powerdecreases) from the center toward the lens periphery. The focal range ofa hyperbolic lens can be rather large and is controllable by so calledconic constant or shape parameter defining the hyperbolic surface shape.

Examples of the distribution of refractive power in a lens withhyperbolic surface is shown in FIG. 2 where local refractive power inDiopters m⁻¹ are plotted against the distance from the optical axis inmm. It is understood that such an optical profile with refractive powerdecreasing with the distance from the optical axis, or with the apertureof the imaging system, or pupil diameter of the eye, can be present inthe originally implanted lens, or it may be created by the hydrogelmodification by a laser after lens implantation.

Based on the present studies it is understood that the positivepolyfocality and its changes in the natural lens assist the eye toaccommodate in several ways:

It projects on the retina simultaneously images of all objects in thefield of view in all distances covered by the focal range of the lens.This significantly (by more than 1 Diopter) increases the depth of thefocus of the eye since all objects create a well-focused image(accompanied by many dis-focused images that the brain learns tosuppress).

The natural lens increases its hyperbolicity due to the accommodation,which further increases the focal range of the lens and, therefore, thedepth of the focus still further.

The eye helps to focus on near objects by narrowing the pupil. This socalled “pupillary reflex” or “near myosis” has two consequences: first,it decreases the aperture and thus increases depth of the focus of theeye as the optical system (narrowing aperture blocks rays that are farfrom the axis and coming in at sharp angles with respect to the axis);and it increases the average refractive power of the lens by using onlyits central portion with the highest refractive power.

It is obvious from our studies that near myosis can assist the nearfocus only for lenses with hyperbolic aberration, i.e. with positivepolyfocality. It has little effect in monofocal parabolic lenses, and itis counterproductive in lenses with negative polyfocality: spherical orelliptical (e.g., meniscoid) lens becomes weaker lens with lowerrefractive power by the near myosis rather than a stronger lens that isneeded for near focus.

An artificial lens according to the invention is a hydrogel deviceimplantable into the posterior chamber of human eye for replacement ofthe natural crystalline lens. It is designed to mimic or replicateessential physiological and optical functions of natural lens withoutcreating problems that earlier attempts could cause in some situations.It is important to recognize that this is achieved by a novel thoughtfulcombination of features that might have been individually, or indifferent combinations, applied previously with a lesser success. Thenatural lens also achieves its function due to its balanced combinationof features rather than to a single feature.

The features contributing to the overall function and combined accordingto the invention include size and shape of the implant; materialproperties; surface properties; optical properties; implantation method;and manufacturing method. We will describe the various features belowand provide exemplary configurations of how individual features mutuallyinteract to provide beneficial effect. It is important to recognize thatthe implant may combine several of the described features to achievedesirable effects, however, the invention is not limited to theexemplary configurations described below and includes variouscombinations of features.

Referring to FIGS. 3A and 3B, the implant has a main optical axis 1Awith a central optical part 2 and a peripheral supporting part 3. Theoverall shape of the implant is defined by its anterior surface 4,posterior surface 5 and the transition surface 6 between the upperboundaries 7A and 7B of the anterior and posterior faces, respectively.Each face is composed of two or more surfaces. The anterior centraloptical surface 8A has boundary 9A and central posterior optical surface8B has boundary 9B. Each of the surfaces may be divided into two or morezones with the boundary between them (denoted 13A and 13B in FIGS. 5A to5C) being circles, straight lines or otherwise defined shapes. Theapexes of the central anterior optical surface 10A and central posterioroptical surface 10B are positioned on the main optical axis 1A. Theanterior peripheral supporting surface is 11A and the posteriorperipheral supporting surface is 11B.

Referring to the FIGS. 6A, 6B and 6C describing lenses comprisingseveral different materials or layers, it is understood that any layeror structure in the optical path may be formed by the hydrogel accordingto the present invention. The preferred layer may be the layer closestto the cornea, i.e. the one forming the anterior optical surface of thelens. Optical structures formed by the hydrogel optical modification maybe refractive structures, diffractive structures, Fresnel lenses shownin the FIG. 6B, refractive index gradient lenses or similarly Suchstructures can be formed within the hydrogel, or on its surface,preferably on the anterior optical surface.

The boundaries 7A and 7B are distinguishable as a discontinuity on thetop of the anterior and posterior surfaces 4 and 5, respectively. Such adiscontinuity lay in the inflexion point of the surface in the directionof the optical axis, or a in a point of discontinuity of the secondderivative of the surface in the direction of the optical axis. Theboundary can be rounded and continuous, but advantageously it is formedby a sharp rim or edge. The advantage of the sharp edge is in formingthe obstacle to migration of cells such as fibroblasts along the capsulesurface (the usual reason for posterior capsule opacification).

The overall lens diameter is defined as the larger diameter of theboundaries 7A and 7B. The lens optical zone diameter is defined as thesmallest diameter of the boundaries 9A and 9B. The posterior sagittaldepth is the vertical distance between the posterior apex 10B and theplane defining the posterior boundary 7B. Central thickness is thedistance between apexes 10A and 10B. Anterior depth is the verticaldistance between the anterior apex 10A and the plane defining theanterior boundary 7A.

The main optical axis 1A may be the axis of symmetry in the case thatboundaries 7A and 7B, as well as boundaries 9A and 9B, are defined bycircles in the plane perpendicular to the optical axis, and if thecentral optical part 2 is symmetrical and e.g., does not have anycylindrical component. Such implant with symmetric circular footprint isshown in FIG. 3B. However, the rims and boundaries may have other thancircular footprint, e.g. elliptical as shown in FIG. 4A, or may have thefootprint shaped as a truncated circle in FIGS. 4B to 4E with single,double, triple or quadruple truncating cuts 12A to 12D. These truncatedfootprint shapes serve several purposes:

They provide better access into the space behind the lens during theimplantation. It is important to clean this space well in order toremove any viscoelastic polymers or lubricants or other auxiliary agentsbefore the surgical incision is closed.

They prevent rotation of the lens after the capsule shrinks around theIOL. This is particularly important for toric lenses.

They facilitate folding and insertion through a small incision.

In the case that the optics has a cylindrical component, then thecylinder axis 1B will be positioned in a defined way with respect to theasymmetry of the outside rim, e.g. be in the angle α to the truncatingcuts 12A and 12B as shown in the FIG. 4F. Needless to say that thetruncating cuts 12A to 12D may not be necessarily straight cuts, but maybe suitably formed to e.g. a crescent shape, and their number may beeven higher than 4. Also, the truncating cuts may not be of the samelength or positioned symmetrically. It can be appreciated that thefootprint with truncated rim will facilitate folding of the implant andits insertion through a small surgical incision. In addition, theasymmetric rim footprint will prevent the implant rotation once thecapsule settles around it. This is particularly important for toriclenses with a cylindrical component designed to compensate forastigmatism.

The posterior surface 5 is shaped and sized to approximate the shape andsize of the posterior surface of the natural lens and to achieve contactwith at least the major part of the posterior capsule of the eye. Thisis important for several reasons:

The implant will keep the posterior capsule in its natural shape,unwrinkled and smooth for the optimum optical performance;

The tight contact between the capsule and the implant will preventmigration of fibroblasts that could cause the posterior capsuleopacification; this is particularly effective if the posterior surfaceis highly hydrated and carrying fixed negative charge.

The implant will occupy the space vacated by the posterior side of thenatural lens and keep thus vitreous body from advancing forward andprevent thus the decrease of the pressure of vitreous body againstretina (which could facilitate retinal detachment and/or cystoic macularedema).

It should be noted that the intimate contact between the implant andposterior capsule is beneficial particularly if the contacting surfaceof the implant is hydrophilic and carrying fixed negative charge inorder to prevent capsular fibrosis and its consequent stiffening,opacification and contraction that would interfere with the implantfunction (or could even dislocate it), as will be described hereinafter.

In the preferred embodiment of the invention, at least the major part ofthe posterior surface 5 is formed by a generally smooth convex surfaceformed by rotation of conic sections around the optical axis, or acombination of such surfaces. The peripheral part is preferably formedby a conic surface or a hyperboloid surface, while the central opticalsurface is preferably hyperboloid, paraboloid or spherical surface (or acombination thereof). The sagittal depth of the posterior surface (i.e.the vertical distance between the posterior central optical surface apex10B and the boundary of the posterior surface 7B, measured on the mainoptical axis 1A) should be larger than 1.1 mm in order for lens toperform its function well. To perform well in the whole refractiverange, the posterior sagittal depth should be larger than 1.25 mm,advantageously larger than 1.75 mm and preferably larger than 2 mm, butin any case less than about 2.75 mm.

The overall outer diameter of the implant (LOD) is important for itscentricity, position stability and capsule-filling capability. The outerdiameter of the posterior surface 5, i.e. the largest dimension of theposterior outer boundary 7B (in the plane perpendicular to the main axis1A) should be larger than 8.4 mm, desirably at least 8.9 mm andpreferably at least 9.2 mm. The largest outer diameter permissible isabout 11 mm, but desirably should be lower than 10.75 mm and preferablyat smaller than 10.5 mm. The considerable flexibility in the outerdimensions is allowed by several factors—flexibility of the lens, andparticularly flexibility of the outer peripheral supporting part 3 thatcan accommodate various capsule sizes and capsule contraction withoutdeforming the central optical part 2.

The central optical surfaces may consist of one or more zones withdifferent geometry. The zones may be concentric, in which case theposterior boundary 13B between them in the FIG. 5A will be circular.Zones may also be divided by straight boundaries, in which case thezones may have crescent or wedge footprint. Various examples are shownin FIGS. 5A to 5C. The zones may be on the anterior or posterior opticalsurface. FIG. 5A shows the posterior optical surface is divided by theboundary 13B into two concentric optical zones—the central optical zone8B1 and the outer optical zone 8B2. For instance, the posterior opticalsurface of the central optical zone 8B1 may be a spherical or paraboliczone used for the sharp near vision, while the hyperbolic outer zoneserves for intermediate and far distance vision. Alternatively, bothzones may have hyperboloid surfaces with different central radii Roand/or different conic constants. Each optical surface may be alsodivided into more than two zones. The example in FIG. 5B shows the topview of the lens which anterior optical surface 8A is divided by astraight boundary 13A into two optical zones of equal area 8A1 and 8A2.Each of those zones has different shape with different opticalparameters. The example in FIG. 5C shows a top view of a lens withanterior optical surface 8A divided by two straight boundaries 13A and13B into four paired optical zones 8A1 and 8A2, each having a differentarea and different optical parameters. For instance, 8A1 may have higherrefractive power that 8A2 and serve for near focus. One of the zones mayhave a cylindrical component.

Both optical surfaces (or their zones or segments) are surfaces formedby rotation of a conical section along the optical axis, or by acombination thereof. One or both optical surfaces may contain one ormore spherical optical zones. Advantageously, at least one of theoptical surfaces comprises at least one hyperbolic surface, preferablyin the outer optical zone. Preferably, both optical surfaces comprise atleast one hyperbolic zone each. Such hyperbolic surface resembles thesurfaces of the NCL and mimics some of its beneficial opticalproperties. Even more preferably, both posterior and anterior opticalsurfaces are hyperbolic surfaces or a combination of two or moreconcentric hyperbolic zones. Lenses with at least one hyperbolic surfacehave so called hyperbolic aberration, the very opposite of sphericalaberration of lenses with spherical, ellipsoid or meniscoid surfaces.The lenses with hyperbolic aberration have highest refraction in thecenter and gradually decreasing with distance from the optical axis. (Inlenses with spherical aberration the refractive power increases withdistance from the optical axis.) The hyperbolic aberration helps the eyeto accommodate through several mechanisms described above. It should beunderstood that hyperbolic aberration can be created not only by theexactly hyperbolic surfaces in the geometry sense, but also by similarsurfaces where surface steepness generally decreases with the distancefrom the optical axis. Therefore, by “hyperbolic surfaces” are meantalso other hyperbole-like surfaces approximating this property.

As the spherical aberration is measured with increasing aperture (orpupil diameter) the negative spherical aberration of hyperboloid-likesurfaces increases in absolute value (i.e., gets more negative). Thespherical aberration can be expressed in various alternative ways, suchas a deviation of wave-front in microns, or as steepness of decrease oflocal refractive power from the optical axis, or by decrease of therefractive power with increasing aperture, or by the value of conicconstant or shape parameter corresponding to such optical profiles.Those skilled in the art can readily convert one of such values into thecorresponding value expressed in a different way. The implant accordingto the present invention can have originally any spherical aberration,since its value can be adjusted post-operatively using the method ofthis invention.

The spherical aberration in the final implanted state (i.e., after thehydrogel modification by a laser) will be generally between about −0.1microns and −2 microns on aperture 4.5 mm. Preferably, the finalspherical aberration will be between about −0.5 microns and −1.5 micronson aperture 4.5 mm, and even more advantageously the sphericalaberration will be between about −0.75 microns and −1.25 microns on theaperture 4.5 mm.

In order to mimic the optical properties of the NCL, conical constantsof the anterior and posterior optical surfaces are selected so that therefractive power of the central optical part 2 generally decreases fromthe highest value at the optical axis to the lowest value at theperiphery of the central optical part 2.

The steepness of the refractive power decrease with the distance fromoptical axis is dependent on the shape parameter (conic constant) of thehyperbolic surface. The conic parameter should be selected that theaverage decrease of the refractive power is between −0.25 Dpt/mm and −3Dpt/mm, advantageously between −0.5 Dpt/mm and −2.5 Dpt/mm andpreferably between about −1 Dpt/mm and −2 Dpt/mm.

The posterior central radius of curvature (at the point where theoptical axis intersects the posterior apex) is advantageously from 2.5to 8 mm, and preferably from about 3.0 to 5 mm. The conic constant ofthe posterior surface is advantageously selected from the range of about+3 to about −14 reported for NCL, preferably from about −1 to −8.

The central radius Ro of the anterior optical surface 8A is selected tobe either larger than about +3 mm or smaller than about −3 mm, andpreferably larger than from about +5 mm or smaller than about −5 mm.

The conical constant of the anterior optical surface 8A is selected fromthe range from +6 to −22 reported from human NCL, preferably from therange between about −1 to −8 mm.

The anterior optical surface 8A may be formed partly or fully by aspherical surface or a parabolic surface. In that case the centralposterior optical surface 8B should be preferably hyperbolic with theconic parameter selected in such a range so that the whole lens hashyperbolic aberration.

Preferably though, at least the major part of the anterior opticalsurface 8A is a hyperboloid surface, particularly the outer opticalzone. The central optical zone of the anterior optical surface havingdiameter between about 1.5 to 4 mm, advantageously between about 2 and3.5 mm, can be formed by parabolic or spherical surface in order tofurther improve the near focus resolution.

FIG. 2 shows schematically one example of the preferred optical profileof the lens according to the invention. It should be appreciated thatdifferent eyes require different refractive power of the implanted lens.

Most of the current IOLs are not bioanalogic since they are designed tosimulate just the basic optical function of NCL, i.e. to provide thebasic refractive power needed to focus a distant object on retina.Depending on the specific eye, the basic refractive power is usuallybetween 15 and 30 Dpt, with some deviations on either side. Thisrequirement can be met by an approximately monofocal (usually spherical)rigid lens located somewhere near of the principal plane of the NCL.Since most detailed images are projected onto a relatively small part ofretina (macula) located on the optical axis, and since many of ouractivities are performed at small eye aperture (constricted pupil), mostIOLs are significantly smaller than the NCL (4.5 to 6 mm for most IOLsas opposed to 9.5 to 10.5 mm for the NCL). The small size of optics ispreferred by some IOL manufacturers for easier adaptation of such IOLfor implantation through a small incision. For the same reason, mostIOLs are made from a soft, elastic material that allows implantationthrough a small incision in a deformed (folded, rolled, etc.) shape.This deformability has no relation to the optical function, however.

Small size of optics has its disadvantages, however. IOL edges mayreflect light at large pupil opening (e.g., during night driving) andcause glare, halos and other adverse effects. Besides, a small opticcannot project all peripheral and off-axis rays that NCL does,particularly at a large pupil opening. Lastly, a small size opticsinterferes with clear visibility of retinal periphery that is sometimesneeded for diagnostics and treatment. For those reasons, the largeoptics similar in size to NCL is preferable over a smaller one that isused in most of the current IOLs. Importantly, the whole large opticalzone has to have well defined geometry to be optically useful. Lenseswith meniscoid optical surfaces have poorly defined shape particularlyin the peripheral region. This may cause unexpected and disturbingoptical phenomena.

Some modern IOLs are designed to simulate to some extent theaccommodation or pseudoaccomodation of the NCL (i.e. allowing the eye tofocus on both far and near objects). Various IOLs use different means toachieve this goal: some are using bifocal, multifocal or polyfocaloptics; others are using designs allowing anterior-posterior shift ofthe IOL optics with respect to the eye; or allow change of optical powerby changing mutual position between two lenses. Some lenses even changethe refractive power due to liquid transfer within the lens driven bypressure of cilliary muscles and/or vitreous body, change of headposition or by a miniature pump.

These designs are sometimes rather intricate contraptions, verydifferent in size, shape and material properties from the NCL. Thismakes them susceptible to various problems, such as fibrosis of thecapsule or cell ingrowth or protein deposits on their surfaces thatinterfere with their function. In addition, their increased bulk andcomplicated design interferes with the need of all modern IOLs to beimplantable through a small incision. This requires designs withsmall-diameter optics and use of materials with high refractive indexthat are more reflective than the NCL, increasing thus the glare andhalo problems.

In most cases, these lenses are using optics of a small diameter,typically 4.5 to 6 mm, with slender, flexible “haptics” to position theoptics in the center of the optical path. In addition, deformablematerials are used to allow folding or rolling for implantation througha small incision. The surface properties of such IOLs are sometimesmodified to achieve better biocompatibility (e.g., A. M Domschke in theUS Pat. Publ. No. 2012/0147323, J. Salamone et al in the US Pat. Publ.No. 2008/0003259).

This common design allows folding the IOL for the implantation through arelatively small incision (usually 2 to 3 mm). However, the small IOLsize has its own drawbacks:

The small optics with diameter 6 mm or less may not fully replace thecrystalline lens of diameter 9 to 10.5 mm if eye aperture is large dueto poor light conditions (causing night glare, halos, limited peripheralvision etc.) or if the IOL becomes decentered (causing the “sunsetsyndrome” or other problems);

Small optics cannot project all peripheral and off-axis rays the NCLdoes, reducing thus the imaging performance particularly at large pupilopenings (needed for e.g., night peripheral vision);

Small optics may complicate or even prevent retinal examination andtreatment (which may be important particularly in the case ofdiabetics).

In addition, a small IOL size leaves essentially vacant the space thatwas originally occupied by the much larger NCL. Consequently, thevitreous body is allowed to advance and its pressure against the retinais partly relieved. This may cause an increased probability of retinaldetachment after the cataract surgery as reported by J. A. Rowe, J. C.Erie, K. H. Baratz et al. (1999). “Retinal detachment in Olmsted County,Minn., 1976 through 1995”. Ophthalmology 106 (1): 154-159. The sameeffect may also cause or facilitate Cystoid Macular Edema (CME). SeeSteven R. Virata, The Retina Center, Lafayette, Ind.: Cystoid MacularEdema, WEB page.

There is another disadvantage of a small optics and the conventional IOLdesign with haptics: The IOL with optics suspended in the relativelyvacant space by means of relatively fragile haptics may be sensitive todamage and/or dislocation in case of an accidental impact (fall on aslippery surface, car collision, a punch, etc.).

Some problems derived from a small bulk of IOLs and small-diameteroptics are being addressed by IOL designs that fill the space vacated byNCL to a smaller or larger extent. There are several approaches to this,each with its own advantages and disadvantages:

Capsule-filling by a liquid that can solidify into a clear, flexiblesolid such as a silicone rubber. As long as the filler material hassimilar deformability as the NCL, it was expected that this approachwould restore the natural lens accommodation (e.g., Gasser et al. inU.S. Pat. No. 5,224,957). However, the materials used so far often causefibrosis and opacification of the capsule. Besides, it is difficult tocontrol the shape and optical parameters of the in situ formed IOL

Implantation of a large, bulky IOL in a highly deformed shape that allowimplantation through a reasonably small incision and fills a significantpart of the capsule. This approach was tried with hydrophobic memorypolymers that can be “frozen” in a highly deformed shape forimplantation, and returns into the original functional shape uponheating to body temperature (Gupta in U.S. Pat. No. 4,834,750 and USPat. No. RE 36,150). However, the hydrophobic memory polymer is veryforeign material and causes similar problems like the materials used tofill the capsule.

Similar approach was also tried with hydrogels. Very large IOLs,mimicking size and shape of the natural lens, have been implanted intothe vacated capsule (e.g., Wichterle '732 and Stoy '283). The problem ofthese particular IOLs was their peculiar optics. These lenses hadmeniscoid anterior optical surfaces that deviated strongly from thegeometry of an NCL. The meniscoid shape was formed by solidification ofa free surface of the monomer mixture, and there was a problem withcontrol of the optical properties of such IOLs. In addition, theselenses were often too bulky for implantation through a small incision.Moreover, some of the hydrogels used in these lenses lacked the fixednegative charge, and such hydrogels have tendency to calcify sometimeafter their implantation. Some other capsule-filling lenses (Sulc et al.'083 and '903) had anterior protrusions touching the iris andstabilizing thus the lens in the approximately central position butcausing various problems such as blockage of the liquid flow,deformation of lens optics and iris erosion.

Another approach was implantation of a hollow lens (or a lens shell)that was filled after implantation by a liquid solidifying in situ(e.g., Nakada, et al. in U.S. Pat. Nos. 5,091,121 and 5,035,710).

Another approach was implantation of dual-optics IOLs with two lenses,one being in contact with anterior and the other with the posteriorcapsule, both lenses being kept apart by flexible members or connectors(U.S. Pat. Nos. 4,946,469; 4,963,148; 5,275,623; 6,423,094; 6,488,708;6,761,737; 6,764,511; 6,767,363; 6,786,934; 6,818,158; 6,846,326;6,858,040; 6,884,261).

Such implants filling essentially the whole capsule of the originalcrystalline lens have also some problems:

Unless made from extremely biocompatible materials with similarhydration and negative charge as NCL, the anterior face of the IOL maytouch the iris and cause its erosion, depigmentation, bleeding orinflammation.

Some materials are made more biocompatible by having high equilibriumwater content. However, that decreases their refractive index far belowthe optimum value (the value for young NCL).

Important for the IOL is not only the shape and optics type, but alsoits material. An NCL is composed of an intricate natural hydrogelstructure comprising water, salts, and polymeric component containingcollagenous proteins, polysaccharides and proteoglycans. Importantly,the polymeric components contain a considerable concentration of acidicionizable groups, such as carboxylates or sulfates. These groups providethe lens material with a fixed negative charge. The hydration and thenegative charge influence the interaction between the NCL and proteinsin the intraocular fluids. Furthermore, its surface properties affectthe interaction between the lens and cells. It is known that synthetichydrogels containing surface with a fixed negative charge do not attractthe proteins and cells and make hydrogel more resistant to calcification(Karel Smetana Jr. et al, “Intraocular biocompatibility ofHydroxyethylmethacrylate and Methacrylic Acid Copolymer/PartiallyHydrolyzed Poly(2-Hydroxyethyl Methacrylate),” Journal of BiomedicalMaterials Research (1987) vol. 21 pp. 1247-1253), and are not recognizedas a foreign body by immune system (Karel Smetana Jr. et al, “TheInfluence of Hydrogel Functional Groups on Cell Behavior”, Journal ofBiomedical Materials Research (1990) vol. 24 pp. 463-470). Although manyIOL manufacturers avoid materials with carboxylate groups based on theassumption that carboxylates attract calcium ions and thereby causecalcifications, there are several references to hydrogel IOLs containingcarboxylate groups (Wichterle '732, Sulc et al. '083 and '903, Stoy inin U.S. Pat. No. 5,939,208, Michalek and Vacik in '093).

Carboxylate groups may be uniformly dispersed in the hydrogel, orconcentrated mainly on the surface forming a gradient of swelling andcharge density, as described e.g. in Stoy '208 and Sulc et al. U.S. Pat.No. 5,158,832. Typically, the NCL material contains, on average, about66% by weight of water. However, the NCL is structured with denser coreand more hydrated jacket and the NCL hydration changes with age and fromindividual to individual. Therefore, one cannot assign a single watercontent value to the NCL other than average.

Similarly, various layers of the NCL have different refractive indices.The refractive index of the lens varies from approximately 1.406 in thecentral layers down to 1.386 in less dense layers of the lens. See e.g.Hecht, Eugene. Optics, 2nd ed. (1987), Addison Wesley, ISBN0-201-11609-X. p. 178. Therefore, the optically meaningful equivalentrefractive index, or ERI, is given as the characteristic of the NCL.Both refractive index and water content change with the lens age.Average ERI=1.441−3.9×10{circumflex over ( )}−4×AGE, decreasing thusfrom about 1.441 at birth to about 1.414 at 70 years. See M. Dubbelmanet al. “The Shape Of The Aging Human Lens: Curvature, EquivalentRefractive Index And The Lens Paradox”, Vision Research 41 (2001)1867-1877, FIG. 9.

In addition, the ERI increases with accommodation by about 0.0013-0.0015per Diopter. See M. Dubbelman et al, “Change In Shape Of The Aging HumanCrystalline Lens With Accommodation”, Vision Research 45 (2005),117-132Ref pp. 127-128. One can speculate that this change of refractiveindex is related to a change (decrease) of water content due to the lensdeformation during the accommodation. Disregarding these complications,we will use the average ERI=1.42 unless stated differently.

Interesting to see that it is very difficult, if not impossible to finda synthetic hydrogel with same water content and—at the sametime—refractive index as the NCL material. Specifically, a synthetichydrogel containing 66% by wt of water would typically have a refractiveindex of about 1.395 rather than 1.42 that would be expected withhydrogel containing closer to 50% of water.

The average liquid contents for ERI=1.441 (very young average NCL) wouldbe 40% of water while for ERI=1.414 (old average NCL) would need ahydrogel with water content about 55% by weight. Since we believe thatfor bioanalogic IOL material it is more important to simulate refractiveindex than water content of NCL, we have selected the desirable averagewater content range of the IOL according to an exemplary embodiment ofthe invention, to be between 40% and 55% by weight. Of course, this isthe average water content—similarly as with the NCL, the lens may havevarious layers with different water contents, e.g. inner parts withhigher refractive index and outer layers with lower refractive index.

A number of prior art references mention IOLs from hydrogels with highwater content, however, they do not recognize the relation between thewater content and the refractive index value. For instance, Wichterle'732 specifies the desirable refractive index value around 1.4 (broadlyfrom 1.37 to 1.45, which is clearly impossible for known synthetichydrogels with the specified water content: at least 60% and preferably65 to 70% translates into the refractive index range from 1.39 to1.405). The examples show formulations with a low content of carboxylategroups.

Sulc et al. '083 and '903 disclose water content at least 70% andadvantageously at least 90% on the surface or its part, and mentions55-70% water content in prior art IOLs. A core with higher and a casingwith lower refractive index are mentioned, and the core may have theform of a Fresnel lens. The gradient of both hydration and refractiveindex is optionally obtained by NaOH treatment that achievesreorganization of the hydrogel covalent network. Example 1 of thisreference shows an IOL with water content 88.5%, Example 2 shows the IOLwith water content 81%, and Example 4 shows the lens with water content91%. No water content is given for Example 3.

Charles Freeman in US Pat. Publ. No. 2009/0023835 describes a hydrogelmaterial with water content lower than 55% and refractive index higherthan 1.41 and the sodium ion flux in the range of about 16 to about 20micro.eq-mm/hr/cm², useful particularly for phakic posterior chamberIOLs. No carboxyl or acidic groups are mentioned, although theirpresence is known to increase the ion diffusion flux through thehydrogel.

Hydrogel character of the NCL material has some possible, less obviousbut potentially important consequences: its water content is dependenton the pressure against the lens. Consequently, the NCL adjusted to thefar distance may have a different water content, and therefore adifferent refractive index, than the relaxed lens adjusted to the nearobjects. Since the stress in the NCL adjusted for far distance is notdistributed evenly, a gradient of swelling and gradient of refractiveindex may result. This will create subtle changes in the opticalproperties, in addition to the polyfocality of the NCL surfaces. Thesesubtle changes may be important for vision, and it will be difficult toreplicate them otherwise rather than by using a hydrogel of similarphysical-chemical and optical properties, as well as geometry similar tothat of an NCL. In particular, the hydrogel of the NCL substitute shouldhave a similar refractive index and capability to change water contentby an external stress that can be reasonably expected to act on an NCL.Therefore, the hydrogel used in a bioanalogic IOL should have ahydraulic flow capability for water.

Therefore, at least the part of the implant contacting the posteriorcapsule is made from a transparent flexible hydrogel materialapproximating the optical, hydrophilic and electrochemical character oftissue forming the natural lens.

The anterior part of the IOL may interfere with, or even block the flowof the vitreous humor causing thus increase of TOP and ultimatelyglaucoma. This design often requires a preemptive iridectomy.

Unless made from extremely biocompatible materials with similarhydration and negative charge as an NCL, the large-area contact betweenthe capsule and artificial materials used in current IOLs sometimescause the capsule opacifications, fibrosis, etc. These problems are nowbeing solved by the bioanalogic intraocular lens according to thisinvention.

The central optical part 2 is made of a deformable, elastic material,such as a hydrogel with equilibrium water content between about 35 and65%, advantageously between about 38% and 55% and preferably betweenabout 40% and 50% (all % are weight percent and equilibrium watercontent is water content in equilibrium with intraocular fluid, unlessstated otherwise).

Deformability of the optical part is important both for the implantationthrough a small incision and for its accommodation function. The opticalpart may be constructed as a hydrogel shell with a core composed from aliquid or a soft gel, as shown in the FIG. 6A. FIG. 6A shows across-sectional view of a lens with the posterior hydrogel jacket 14,the softer core 15 and the anterior shell 16. The posterior hydrogeljacket 14 is advantageously integral with the peripheral supporting part3 of the lens and contains the fixed negative charge at least on itsposterior surface. The core 15 can be advantageously made from ahydrophobic liquid, such as mineral oil or silicone oil, or from a softsilicone or acrylic slightly cross linked gel that can be easilydesigned and created by those skilled in the art. Alternatively, thecore can be made or a hydrophilic fluid or a soft hydrogel. The anteriorshell 16 can be made from the same or different material as theposterior hydrogel jacket 14.

In one embodiment, the hydrogel jacket and the soft core 15 haveessentially the same refractive index so that the major part of therefraction takes place on the outer optical surfaces of the lens ratherthan on its internal interfaces. This can be achieved e.g. by making thecore from a silicone liquid or a silicone gel having refractive indexaround 1.42, and making the jacket from a hydrogel with water contentbetween about 41 and 45% of water. By formulating the hydrogel correctlyone skilled in the art can adjust the water content in the hydrogel toachieve the substantial match of the refractive indices. Alternatively,the core and the jacket can have different refractive indices so thatpart of the refraction takes place on the internal interfaces betweenmaterials.

FIG. 6B a cross-sectional view of a lens with an internal interfacebetween the core 15 and adjacent optical medium 16 that is shaped toform a compound lens, e.g. a Fresnel lens. The materials of core 15 andthe optical medium 16 have different refractive indices, and one of themis advantageously a fluid that can improve both deformability andrefraction. The zone 15 or 16 (the one with the lower refractive index)can be created by the modification of a hydrogel according to thepresent invention using a laser. The hydrogel modification can becarried out either preoperatively or postoperatively. Advantage of thisarrangement is the possibility to use hydrogels with high water contentand low refractive index as the basic construction material, and yetachieve relatively low central thickness of the lens that allowsimplantation through a small incision.

FIG. 6C shows an alternative design of the lens comprising two differentmaterials. Material on the posterior side 14 is a hydrogel with highhydration rate and containing negatively charged groups. It is the samefor the optical and supporting part. The anterior side material of core15 is a material with lower water content and higher refractive index.The interface between the two materials is refractive.

Both central optical anterior surface 8A and central posterior opticalsurface 8B have a diameter larger than about 5.6 mm, advantageouslylarger than about 6.5 mm and preferably larger than about 7.2 mm.Optimum diameter of the larger of the two optical surfaces is largerthan about 7.5 mm, advantageously about 8 mm to approximate the size ofthe NCL optics. Such a large optic is usually suitable forconvex-concave or plano-convex central optical part 2. For a biconvexoptical part, the anterior optical diameter is usually selected smallerin order to minimize the central thickness of the optical part. In anycase, the diameter of the anterior optical surface 8A is advantageouslynot larger than the diameter of the central posterior optical surface8B.

The central optical surfaces 8A and 8B are surrounded by boundaries 9Aand 9B that are not necessarily circular. The boundaries 9A and/or 9Bmay be also elliptical or have a shape of a truncated circle, in orderto facilitate the lens folding and implantation through a smallincision. Non-circular optical surfaces are particularly suitable forlenses with a cylindrical component.

The posterior peripheral supporting surface 11B is formed by a convexsurface, advantageously a hyperbolic or conical surface with the axisidentical with the main optical axis 1A. This surface is highlyhydrophilic and carrying a fixed negative charge due to a content ofacidic groups such as carboxylate, sulfo, sulphate or phosphate groups.This combination of hydration and negative charge prevents a permanentadhesion to the capsule, prevents migration of cells, particularlyfibroblasts, along the interface between the lens and the capsule,decreases irreversible protein adsorption, and discourages capsularfibrosis and opacification. The posterior peripheral surface isadvantageously limited by a sharp edge 7B that further discourages cellmigration toward the optical zone.

The anterior peripheral supporting surface 11A is a concave surface withits apex located on the optical axis and it is preferably symmetricalalong the axis 1A. Advantageously it is a conical or hyperbolic surfacewith its axis coinciding with the main optical axis 1A. The surface isadvantageously highly hydrophilic and carrying fixed negative charge inorder to discourage cell adhesion and migration and anterior capsularfibrosis. The anterior peripheral surface is advantageously limited by asharp edge 7A that further discourages cell migration.

The anterior and posterior peripheral supporting surfaces 11A and 11Btogether with the connecting surface 6 define the shape of theperipheral supporting part 3. The peripheral supporting part is convexon the posterior side and concave on the anterior side, the averagedistance between the two surfaces ranging from about 0.05 to 1 mm,advantageously from about 0.1 to 0.6 mm and preferably from about 0.15to 0.35 mm. The optimum distance depends on the stiffness of thematerial that is dependent on water content, negative charge density,crosslinking density and other parameters.

If the posterior and anterior surfaces are formed by surfaces of similargeometry, such as hyperbolic surfaces, then the peripheral supportingpart 3 will have even thickness. The arrangement shown in FIG. 7A hasthe advantage to be readily deformable and adjustable to various sizesof the capsule, and two sharp edges 7A and 7B preventing migration offibroblasts toward the optical zone.

The peripheral supporting part 3 can be also made less or moredeformable by increasing or decreasing its thickness from the rim towardthe center, as shown in FIGS. 7B and 7C, respectively. These figuresalso show various alternative arrangements of edges 7A and 7B.

The anterior surface 4 of the implant is shaped to avoid any permanentcontact with iris that could cause iris erosion, pupilar block, irispigment transfer to the implant and other problems. Such a contact couldalso interfere with the flow of the intraocular fluid causing thusadverse changes of the intraocular pressure. It could also interferewith the contraction of the pupil as to prevent so called near myosisthat helps the near focus both by the natural lens and by the implantaccording to the invention. Therefore, the anterior central opticalsurface 8A part is partially sunk due to the anterior peripheralsupporting surface 11A concavity and due to positioning the boundary 9Aunder the plane defined by the anterior boundary 7A. The centralanterior surface 8A is a plane, a convex surface or a concave surfacewith its anterior apex 10A not exceeding the uppermost point of the lens(the higher of 7A and 7B) by more than about 0.25 mm, advantageously notexceeding the upper rim at all and preferably having the anterior apex10A bellow the uppermost point 7A by at least 0.1 mm.

At least the major part (including the central optical surfaces 8A and8B) of both anterior and posterior surfaces 4 and 5 are defined byrotation of one or more conic sections around the main optical axis 1A.wherein the term “conic section” includes a segment of a line forpurpose of this application. The surfaces defined by the rotation willinclude a plane perpendicular to the axis and conical surfacesymmetrical by the main optical axis 1A. The peripheral supporting partis convex on the posterior side and concave on the anterior side, theaverage distance between the two surfaces ranging from about 0.05 to 1mm, advantageously from about 0.1 to 0.6 mm and preferably from about0.15 to 0.35 mm.

In at least one embodiment, the lens according to the invention ismanufactured by solidification of liquid polymer precursors. In thepreferred embodiment, the solidification takes place in contact with asolid mold, particularly a mold made of a hydrophobic plastic. It can beappreciated that the surface microstructure of a polymer depends on theenvironment in which its solidification took place. The surfacemicrostructure will be different if the solidification occurs on thesolid liquid interface that if it takes place on the liquid-liquid orliquid-gas interface. Preferably, at least all optical surfaces arecreated by solidification of the precursor on a solid interface. Evenmore preferably, whole surface of the implant is formed bysolidification of a liquid precursor against a solid surface,particularly a hydrophobic plastic surface. Preferred plastic for themold is a polyolefin, and particularly preferred plastic ispolypropylene. The polyolefin has low polarity and has low interactionwith highly polar monomers that are used as hydrogel precursors.Likewise, the hydrogel formed by the liquid precursor solidification hasvery low adhesion to the mold surface and can be cleanly detachedwithout even a microscopic surface damage. This is important for bothoptical properties and for long-term biocompatibility of the implant.

Manufacturing a relatively large lens of a precise shape by molding isdifficult. It is recognized by those skilled in the art that anysolidification of the liquid precursor is accompanied by the volumeshrinkage that may even exceed 20 percent. In a closed mold of aconstant volume, such a shrinkage will prevent copying of the internalmold surface and cause formation of vacuoles, bubbles, surfacedeformities and other imperfections. This is the main reason why themeniscus casting methods described above were used for IOL molding.Other inventors have described a method and a mold design allowing theexcess of monomers to be transported from adjacent spaces by the suctioncreated by the volume contraction (Shepherd T., U.S. Pat. No.4,815,690). However, this method cannot be used in cases where theliquid precursor gellifies at a low conversion (e.g., 5 to 10 percent)due to the crosslinking polymerization.

We have discovered a different method for the volume shrinkagecompensation, namely, decrease of the internal mold cavity volume due tothe deformation of certain mold parts. The mold depicted in FIG. 8 iscomposed from two parts 18A and 18B, the part 18A being used for moldingthe anterior surface 4 and the part 18B for molding the posteriorsurface 5.

The shaping surface 19B of the part 18B has a shape needed to form theposterior optical surface 8B of the lens. The peripheral part 22B of themolding surface has a diameter larger than the diameter of the lens andadvantageously a hyperboloid or conical shape

The part 18A has the shaping surface 19A that is divided into thecentral part 21A shaping the anterior optical surface 8A of the lens,and the peripheral part 22A of the diameter larger than the diameter ofthe lens. The peripheral part 22A has advantageously a hyperboloid orconical shape. The peripheral surface 22A is substantially parallel tothe corresponding surface 22B of the part 18B.

The diameter of the molding the mold parts 18A and 18B are larger thandiameter of the lens and advantageously they are the same. One of thesurfaces for 22A or 22B is equipped with a relatively thin anddeformable barrier 20 with inner surface corresponding to the geometryof the surface 6 of the lens. The height of the part 20 is typicallybetween about 0.05 mm and 1.3 mm, and its thickness is lesser than itsheight. The profile of the part 20 is advantageously wedge-like ortriangular. At least one of its surfaces is advantageously parallel tothe optical axis 1A. The barrier 20 may be separate from the parts 18Aand 18B, but advantageously it is an integral part of one of them.Advantageously, this part 20 is located on the concave surface 22B. In apreferred mode of the operation, the liquid precursor is filled into theconcave mold part 18B in a slight excess to reach over the barrier 20,and then it is covered with the part 18A. The mold is constructed insuch a way that the only contact between parts 18A and 18B is via thepart 20. The solidification of the precursor generates its contractionand the consequent decrease of the pressure in the mold cavity. At a lowconversion, the additional liquid precursor is pulled into the moldcavity. Once the gel-point is reached due to the crosslinking, theprecursor cannot flow anymore. The decreased pressure will causedeformation of the part 20 and decrease of the distance between parts18A and 18B and the consequent decrease of the molding cavity volume.The two-part mold for the IOL according to the invention is preferablymade by injection molding from a polyolefin, advantageously frompolypropylene.

The preferred liquid precursor for the invention is a mixture of acrylicand/or methacrylic monomers with crosslinkers, initiators and othercomponents known well to those skilled in the art. The preferredprecursor composition comprises a mixture of acrylic and/or methacrylicmonoesters and diesters of glycols where monoesters are hydrophiliccomponents and diesters are crosslinkers. The preferred precursor alsocomprises acrylic and/or methacrylic acid or its salts. Itadvantageously comprises also a UV absorbing molecule with apolymerizable double bond, such as methacryloyloxybenzophenone (MOBP).Other possible derivatives of acrylic or methacrylic acid are theiresters, amides, amidines and salts.

Also part of the hydrogel structure are ionizable groups bearing anegative charge, such as carboxylate, sulfate, phosphate or sulfonatependant groups. They may be introduced by copolymerization withappropriate monomers bearing such groups, such as methacrylic or acrylicacid. In this case, the ionogenic functionality will be uniformlydispersed in the hydrogel. Particularly advantageous are hydrogels withionogenic groups concentrated mainly on the surface with the consequentgradient of swelling and charge density. Such gradients can be createdby after-treatment of molded lenses, e.g. by methods described in Stoy'208 and Sulc et al. U.S. Pat. Nos. 5,080,683 and 5,158,832.

Other methods include, e.g. grafting of monomers comprising ionogenicgroups on the lens surface. It is understood that only a part of thelens surface may be treated to contain high concentration of ionogenicgroups, or that different parts of the surface may be treated bydifferent methods.

The lens according to the invention can be implanted in the deformed andpartly dehydrated state. The controlled partial dehydration can beachieved by contacting lens with a suitably hypertonic aqueous solutionof physiologically acceptable salts, such as chlorides, sulfates orphosphates magnesium or monovalent ions, such as sodium or potassium.Salt concentration can be adjusted to achieve hydration between about15% and 25% by weight of the liquid. The lens in the hypertonic solutioncan be advantageously sterilized by autoclaving.

Another method for preparing the hydrogel lens for implantation throughan incision with reduced size is plastification of the hydrogel by anon-toxic organic water-miscible solvent, such as glycerol ordimethylsulfoxide, in such a way that the plasticized hydrogel hassoftening temperature above ambient but lower than eye temperature. Suchcomposition and process is described e.g. in Sulc et al, U.S. Pat. No.4,834,753 that is hereby incorporated by this reference.

The lens according to at least one embodiment of the invention isadvantageously implanted in the state of the osmotic non-equilibrium toadhere to the tissue temporarily. The osmotic non-equilibrium allows thelens centering by adhering it against the posterior capsule while thecapsule shrinks around it. Once the lens is enveloped by the capsule,its position is stabilized. The osmotic non-equilibrium can be achievedin various ways: soaking the lens prior to the implantation in ahypertonic salt solution, e.g. in a solution of 10% to 22% by wt. NaCl,advantageously 15% to 19% by wt.; replacing water prior to theimplantation by a smaller concentration of a water-miscible solvent,such as glycerol or dimethylsulfoxide; or implanting the lens in thestate in which the iogenic groups are not fully ionized, i.e. in theacidic state prior to the neutralization, and letting the neutralizationproceed spontaneously in situ by positive ions from the body fluids. Thelens achieves its osmotic equilibrium spontaneously in hours to daysafter the implantation.

The lens shape is being formed preferably by crosslinkingcopolymerization of methacrylic and/or acrylic esters and salts in theclosed two-part mold.

The shape of the lens can be adjusted after the molding by removing somepart of the lens, e.g. by cutting off part of the supporting part, bydrilling the lens outside the optical zone etc. The shape adjustment canbe made in the hydrogel or the xerogel (i.e. non-hydrated) state. Wehave found that the negatively charged hydrogel material even allows useof methods developed primarily for living tissues (incl. NCR), such asultrasonic phacoemulsification, cauterization or femtosecond lasertreatment. These methods allow shape adjustment even in the fullyhydrated hydrogel state. The femtosecond laser may be used even forformation of cavities inside the hydrogel lens that can be used to formnew refractive members in the lens, for instance as a refractivecylindrical lens for astigmatism compensation. In the case that thematter removed by the shape adjustment (e.g., by a laser treatment) iswater-soluble and substantially non-toxic, such an optical adjustmentcan be conceivably achieved even post-operatively in situ. Thecomposition of the hydrogel in at least the treated part of the lensshould be advantageously based on esters of polymethacrylic acid. It isknown that such polymers are capable of depolymerization to their parentmonomers (such as 2-hydroxyethyl methacrylate or methacrylic acid) thatare well soluble, easily diffusible compounds of low toxicity. Otherpolymers, such as polyacrylates, polyvinyl compounds or polyurethanes donot have this advantage.

The invention is further illustrated by the following Examples that aremeant to provide additional information without limiting the scope ofthe invention.

Example 1

The following monomer mixture was prepared: 98 weight parts of2-hydroxyethyl methacrylate (HEMA), 0.5 wt % of triethyleneglycoldimethacrylate (TEGDMA), 1 wt % of methacryloyloxybenzophenone (MOBP), 1wt % of methacrylic acid, 0.25 wt % of camphorcquinone (CQ) and 0.05 wt% of trieathanolamine (TEA). The mixture was de-aired using by carbondioxide and filled into two-part plastic molds shown schematically inFIG. 8 where 18B is the part of the mold for molding the a posteriorlens surface, 18A is the part of the mold to shape the anterior part ofthe surface of the lens. Both parts are injection molded frompolypropylene (PP). The shaping surface 19B of the part 18B has shapeformed by two concentric hyperboloids. The central part of the surfacehas the diameter 3 mm, central radius of 3.25 mm and conic constant−3.76 while the peripheral is hyperboloid with central radius of 3.25 mmand conic constant −6.26. The molding surface is equipped with aprotruding circular barrier 20 on diameter 8.5 mm that has asymmetrictriangular profile, height 0.2 mm. This lip is designed to shape theconnecting surface 6 in FIG. 3A.

The part 18A has the shaping surface 19A that is divided into thecentral part 21 of diameter 6.8 mm and the peripheral part 22A of thediameter 13 mm. The peripheral part is formed by a hyperboloid with thecentral radius 3.25 mm and the conic constant −6.26. The peripheralhyperbolic surface is parallel to the corresponding surface of the part18B. The central portion of the part 18A has the central radius ofcurvature −20 mm and conic constant h=1.

About 0.1 ml of the monomer mixture is pipetted into the part 18B, thenit is covered by the part 18A that is carefully centered and pressedgently against it by a small weight. The only direct contact between theparts is the circular contact between the barrier 20 and the peripheralpart of 22A. The mold is then illuminated for 10 minutes by a blue lightat the wavelength 471 nm. The light initiates polymerization of themonomers accompanied by gelling at a relatively low conversion and byvolume contraction that is roughly proportionate to conversion. Thecontraction of the soft gel creates a mild vacuum that pulls both partsof the mold together. The conical peripheral part 22A of the mold 18Apresses against the barrier 20, deforms it slightly and closes to thepart 18B to reduce the volume of the molding cavity. This compensatesfor the volume shrinkage due to the polymerization. The described molddesign is particularly suitable for production of relatively bulky IOLsfrom materials with high polymerization contraction that achievesgel-point at a relatively low conversion.

The mold parts are separated and the xerogel lens, the exact copy of themold cavity, is neutralized by solution of sodium bicarbonate andextracted with isotonic solution. The linear expansion factor betweenthe xerogel and hydrogel lens is 1.17. After evaluation of opticalproperties the lens was immersed in the 18% by weight aqueous solutionof NaCl in a sealed blister package and sterilized by autoclaving.

Example 2

The following monomer mixture was prepared: 94 weight parts of2-hydroxyethyl methacrylate (HEMA), 0.5 wt % of triethyleneglycoldimethacrylate (TEGDMA), 4.5 wt % of methacryloyloxybenzophenone (MOBP),1 wt % of methacrylic acid and 0.25 wt % of dibenzoylperoxide. Themixture was de-aired using nitrogen carbon and filled into two-partplastic molds shown schematically in FIG. 8. The shaping surface 19B ofthe part 18B has a shape formed by two concentric surfaces. The centralpart of the surface has the diameter 3 mm, central radius of 3.00 mm andconic constant 1 while the peripheral section is a hyperboloid withcentral radius of 3.25 mm and conic constant −6.26. The molding surfaceis equipped with a protruding circular barrier 20 on diameter 8.8 mmthat has asymmetric triangular profile, height 0.15 mm. The inner sideof the barrier 20 is designed to shape the connecting surface 6 in FIG.3A.

The part 18A has the shaping surface 19A that is divided into thecentral part 21 of diameter 7.1 mm and the peripheral part 22A of thediameter 13 mm. The peripheral part is formed by a hyperboloid with thecentral radius 3.25 mm and the conic constant −6.26. The peripheralhyperbolic surface is parallel to the corresponding surface of the part18B. The central portion of the part 18A is a plane perpendicular to theoptical axis 1A.

About 0.1 ml of the monomer mixture is pipetted into the part 18B, thenit is covered by the part 18A that is carefully centered and pressedgently against it by a small weight. The only direct contact between theparts is the circular contact between the barrier 20 and the peripheralpart of 22A. The mold is then heated to 75° C. for 6 hours.

The mold parts are separated and the xerogel lens, the exact copy of themold cavity, is neutralized by solution of sodium bicarbonate andextracted 3 times with ethyl alcohol and 5 times with isotonic solution.The lens was yellow with complete absorption of UV light and part of theblue visible light. The linear expansion factor between the xerogel andhydrogel lens is 1.13. After evaluation of optical properties the lenswas immersed in the 15% by weight aqueous solution of NaCl in a sealedblister package and sterilized by autoclaving.

Example 3

The following monomer mixture was prepared: 94.5 weight parts of2-hydroxyethyl methacrylate (HEMA), 0.5 wt % of triethyleneglycoldimethacrylate (TEGDMA), 5 wt % of methacryloyloxybenzophenone (MOBP)and 0.25 wt % of dibenzoylperoxide. The mixture was de-aired usingnitrogen carbon and filled into two-part plastic molds shownschematically in FIG. 8. The shaping surface 19B of the part 18B has ashape formed by two concentric surfaces. The central part of the surfacehas the diameter 6.5 mm, central radius of 4.5 mm and conic constant 0while the peripheral section is a hyperboloid with central radius of4.25 mm and conic constant −8. The molding surface is equipped with aprotruding circular barrier 20 on diameter 9.3 mm that has asymmetrictriangular profile, height 0.35 mm. The inner side of the barrier 20 isdesigned to shape the connecting surface 6 in FIG. 3A.

The part 18A has the shaping surface 19A that is divided into thecentral part 21 of diameter 6.4 mm and the peripheral part 22A of thediameter 13 mm. The peripheral part is formed by a hyperboloid with thecentral radius 4.25 mm and the conic constant −8. The peripheralhyperbolic surface is parallel to the corresponding surface of the part18B. The central portion of the part 18A is a surface of diameter 6.4mm, central radius −3.75 mm and conic constant −6.

About 0.1 ml of the monomer mixture is pipetted into the part 18B, thenit is covered by the part 18A that is carefully centered and pressedgently against it by a small weight. The only direct contact between theparts is the circular contact between the barrier 20 and the peripheralpart of 22A. The mold is then heated to 75° C. for 6 hours.

The mold parts are separated and the xerogel lens, the exact copy of themold cavity, is extracted. The lens is then treated by a quaternary baseas described in the reference Stoy '208.

The z lens from the clear, electroneutral crosslinked hydrophilicpolymer has a surface created by a gradiented layer with high hydrationand negative charge density. The lens was neutralized by solution ofsodium bicarbonate and extracted 3 times with ethyl alcohol and 5 timeswith isotonic solution. The lens was clear with complete absorption ofUV light. The linear expansion factor between the xerogel and hydrogellens is about 1.12. After evaluation of optical properties the lens wasimmersed in the isotonic aqueous solution of NaCl in a sealed blisterpackage and sterilized by autoclaving.

Example 4

Raman spectra of the hydrogels of the invention show a significantdifference between the original hydrogel and the same hydrogel subjectedto TPA by exposure to the focused beam of a femtosecond laser. Namely,there is a significant difference in the ratio of the signal at 3420cm⁻¹ corresponding to water, and at 2945 cm⁻¹ corresponding to the CH₂group from the polymer backbone (see FIG. 9A). The ratio between theintensities of the two peaks is proportional to the phase-shift (innumber of wavelengths) of the test laser beam between the modified andunmodified material (see FIG. 9B). The Raman scan across the modifiedstrip within the hydrogel showed increased water content in the treatedarea (see FIG. 9C). On the other hand, the Raman spectrum of the regionbelow 2000 cm⁻¹ showed no indication of new chemical groups. This isconsistent with replacement of part of the polymer mass for an aqueousliquid through a mechanism such as depolymerization.

These and other advantages of the present invention will be apparent tothose skilled in the art from the foregoing specification. Accordingly,it will be recognized by those skilled in the art that changes ormodifications may be made to the above-described embodiments withoutdeparting from the broad inventive concepts of the invention. It shouldtherefore be understood that this invention is not limited to theparticular embodiments described herein, but is intended to include allchanges and modifications that are within the scope and spirit of theinvention as defined in the claims.

What is claimed is:
 1. An in situ-adjustable ophthalmic implantcomprising a hydrogel wherein said hydrogel comprises an acrylate or amethacrylate copolymer hydrogel wherein said copolymer comprises atleast four co-monomers: a) an acrylate ester or a methacrylate ester,each containing at least one pendant hydroxyl group; b) a polyolacrylate ester, a polyol acrylate amide, a polyol methacrylate ester, ora polyol methacrylate amide, each with at least 2 acrylate ormethacrylate groups per polyol ester or amide; c) acrylic or methacrylicacid, or a derivative thereof, each having at least one pendant carboxylgroup; and d) a vinyl monomer, an acrylic monomer, or a methacrylicmonomer, each having a pendant UV-absorbing group; wherein refractiveproperties of said implant are adjustable by absorption of femtosecondlaser pulses by said hydrogel resulting in a negative change ofrefractive index in selected locations of said implant.
 2. Theophthalmic implant according to claim 1, wherein the implant hasanterior and posterior refractive surfaces forming a lens with positiveor negative refractive power.
 3. The ophthalmic implant according toclaim 2, wherein the lens is an intrastromal lens, an anterior chamberlens, a phakic lens for placement between the iris and a naturalcrystalline lens, or a posterior chamber lens for at least partialreplacement of a natural crystalline lens.
 4. The ophthalmic implantaccording to claim 1, wherein the monomer containing the pendantcarboxyl group is a neutralized or a partially neutralized methacrylicacid.
 5. The ophthalmic implant according to claim 1, wherein themonomer containing the pendant carboxyl group is present in aconcentration between 0.1 molar % and 5 molar % based on all monomerunits of the copolymer.
 6. The ophthalmic implant according to claim 5,wherein the monomer containing the pendant carboxyl group is present ina concentration between 0.5 molar % and 2 molar % based on all monomerunits of the copolymer.
 7. The ophthalmic implant according to claim 1,wherein the monomer containing the pendant UV-absorbing group is presentin a concentration between 0.1 molar % and 5 molar % based on allmonomer units of the copolymer.
 8. The ophthalmic implant according toclaim 7, wherein the monomer containing the pendant UV-absorbing groupis present in a concentration between 0.2 molar % and 2.5 molar % basedon all monomer units of the copolymer.
 9. The ophthalmic implantaccording to claim 1, wherein the pendant carboxyl groups and thependant UV-absorbing groups are present in a molar ratio between 0.25:1and 5:1.
 10. The ophthalmic implant according to claim 9, wherein thependant carboxyl groups and the pendant UV-absorbing groups are presentin a molar ratio between 0.5:1 and 3.5:1.
 11. The ophthalmic implantaccording to claim 1, wherein the pendant carboxyl group is ionized, andthe molar ratio of ionized pendant carboxyl groups to UV-absorbingpendant groups is from 0.5:1 to 3.5:1.
 12. The ophthalmic implantaccording to claim 1, wherein the pendant UV-absorbing group contains aphenolic hydroxyl group conjugated to an aromatic group.
 13. Theophthalmic implant according to claim 1, wherein at least one of thependant UV-absorbing groups is selected from the group consisting ofderivatives of benzophenone, derivatives of benzotriazole, derivativesof coumarin and derivatives of fluorescein.
 14. The ophthalmic implantaccording to claim 1, wherein the copolymer contains at least twodifferent co-monomers containing different pendant UV-absorbing groups.15. The ophthalmic implant according to claim 14, wherein at least oneof the pendant UV-absorbing groups is a benzophenone or a derivativethereof.
 16. The ophthalmic implant according to claim 1, wherein atleast a major portion of the polymer of said hydrogel is a hydrophilicderivative of methacrylic acid.
 17. The ophthalmic implant according toclaim 16 wherein at least a major portion of the hydrophilic methacrylicacid derivative is a glycol ester of methacrylic acid.
 18. Theophthalmic implant according to claim 1, wherein the hydrogel containsmore than 30% by weight of liquid under equilibrium physiologicalconditions.
 19. The ophthalmic implant according to claim 1, wherein thehydrogel contains less than 55% by weight of liquid under equilibriumphysiological conditions.
 20. The ophthalmic implant according to claim1, wherein the hydrogel contains between 35% and 47.5% by weight ofliquid under equilibrium physiological conditions.
 21. The hydrogelophthalmic implant according to claim 1, wherein the hydrogel contains25% to 55% by weight of liquid under equilibrium physiologicalconditions.
 22. The hydrogel ophthalmic implant according to claim 1,wherein the hydrogel contains 25% to 30% by weight of liquid underequilibrium physiological conditions.